Mechanical circulatory support device with centrifugal impeller designed for implantation in the descending aorta

ABSTRACT

Mechanical circulatory supports configured to operate in series with the native heart are disclosed. In an embodiment, a centrifugal pump is used. In an embodiment, inlet and outlet ports are connected into the aorta and blood flow is diverted through a lumen and a centrifugal pump between the inlet and outlet ports. The supports may create a pressure rise between about 40-80 mmHg, and maintain a flow rate of about 5 L/min. The support may be configured to be inserted in a collinear manner with the descending aorta. The support may be optimized to replicate naturally occurring vortex formation within the aorta. Diffusers of different dimensions and configurations, such as helical configuration, and/or the orientation of installation may be used to optimize vortex formation. The support may use an impeller which is electromagnetically suspended, stabilized, and rotated to pump blood.

INCORPORATION BY REFERENCE

This application is a continuation-in-part of U.S. patent applicationSer. No. 14/440,848, which is a U.S. national phase ofPCT/GB2013/052889, filed Nov. 5, 2013, which claims priority to GBapplication No. 1219958.4, filed Nov. 6, 2012, which references areincorporated herein by reference in its entirety for all purposes. Thisapplication also claims priority benefit of U.S. Provisional PatentApplication No. 62/403,428, filed Oct. 3, 2016, and U.S. ProvisionalPatent Application No. 62/513,927, filed Jun. 1, 2017, each of which isincorporated herein by reference in its entirety for all purposes. Anyand all applications related thereto by way of priority thereto ortherefrom are hereby incorporated by reference in their entirety.

BACKGROUND

The present invention relates to a mechanical circulatory support (MCS),otherwise known as a mechanical circulatory support device (MCSD), forassisting or replacing native heart function in cases of congestiveheart failure (CHF).

Patients with CHF usually have a low cardiac output state as the nativeheart functions (pumps) poorly. This in turn leads to poor organperfusion and the symptoms of heart failure including fatigue,breathlessness and feeling generally unwell. In heart failure thekidneys also suffer with poor perfusion and their function oftendeteriorates considerably (a condition called “the cardio-renalsyndrome”). Poor kidney function means that patients feel more unwell,and important drugs have to be withdrawn as they can further adverselyaffect kidney function.

CHF is common and is a significant health care burden. It is graded fromstage I-IV in severity. Once diagnosed a patient has 4-5 years ofprogression from stage I to IV and death. Stage IV patients arebreathless at rest, candidates for heart transplantation, and medicationis considered palliative. Congestive heart failure (CHF) is the maincause of mortality for men and women alike in the western world,affecting about 2% of the population. In the USA alone there are 5.7million patients suffering from CHF and costs to treat this exceed $37.2billion/year. In the Western world current supply of donor hearts onlymeets about 12% of demand. This percentage is higher than the actualnumber because most potential recipients are not included in thecalculation; they are considered not suitable for a transplant becauseof co-morbidities or lack of a matched donor. This shortfall hasresulted in the development of MCS devices as a transplant alternative.MCS devices are expensive and require invasive cardiac surgery(sternotomy or thoracotomy). Implantation carries a significant risk.Not all candidates are suitable for MCS because of co-morbidities.

Most permanent MCS devices assist the ventricle and are attached to itin use. These are called Ventricular Assist Devices (VADs), and aredesigned to drive a flow of blood that is in parallel with flow withinthe native heart, between the ventricle and the aorta. In other words,they are designed as left (or right) ventricular assist devices (LVADsor RVADs), pumping devices that directly unload the respectiveventricle. Such “in-parallel” configurations involve the device andheart sharing, and therefore competing, for inlet flow, which candisrupt normal functioning of the heart. Regeneration of heart musclemay be impeded and the heart is not able to pump to its best capacity.The inlet of most of these VADs is anastomosed to the apex of the leftventricle of the heart, and therefore their installation requires majorsternotomy or thoracotomy and cardiopulmonary bypass (CPB), i.e.stopping of the heart during a prolonged surgical operation, forpermanent installation. Survival rates of patients on VADs have beenpoor.

Due to inefficiencies, existing MCS/VAD devices typically requiresignificantly more input power than is necessary from a theoreticalpoint of view purely to impart the desired momentum to the blood. Theexcess power is used to overcome the losses. The portion of the powerthat is used to overcome flow losses is imparted as unnecessary damageto the blood, leading to increased levels of haemolysis and/or thrombusformation that would be avoided with devices having higher fluid dynamicefficiency.

VADs entered clinical use as displacement (or pulsatile flow) devices,which mimic the native left ventricle by providing pulsatile flow takingover the function of the patient's own left ventricle. Most widely useddisplacement, pulsatile, devices have been extracorporeal devices suchas the BVS® 5000 VAD of Abiomed, Inc. (Danvers, Mass., USA) and theThoratec VAD of Thoratec Corporation (Pleasanton, Calif., USA), andintracoporeal devices such as the Novacor® LVA System of WorldHeart,Inc. (Oakland, Calif., USA), the HeartMate IP and VE/XVE of ThoratecCorporation. Although the large external pneumatic consoles of thefirst-generation displacement VADs have been replaced by implantableelectric systems with a portable controller and power source, theserious problems of device weight (e.g., approximately 1.5 kg for theHeartMate XVE), size, noise, driveline infection and thromboembolismpersist. Consequently, newer displacement devices are totallyimplantable, such as the LionHeart™ VAD of Arrow International, Inc.(Reading, Pa., USA), and the Novacor® LVA System of WorldHeart, Inc.(Oakland, Calif., USA).

Rotary (or continuous flow) devices (second-generation VADs) have beendeveloped to overcome the shortcomings of pulsatile devices. Initialconcerns with their pulseless flow are now overcome, provided that thepatient's native system still provides some pulsatility, and they havetheir own relative advantages (e.g., fewer moving parts, lower powerrequired, absence of bioprosthetic valves) and disadvantages (e.g.,complex control, high afterload and low preload sensitivity, andhaemolysis and thrombosis from unnatural flow patterns). Examples ofaxial rotary pumps (which operate at 10,000-20,000 rpm) are the DeBakeyVAD® of MicroMed Cardiovascular, Inc. (Houston, Tex., USA), theFlowMaker® of Jarvik Heart, Inc. (New York, N.Y., USA), formerly knownas Jarvik 2000, the HeartMate II of Thoratec Corporation (Pleasanton,Calif., USA), and the Impella Recover® system of Impella CardioSystemsAG (Aachen, Germany) intended for short-term circulatory support for upto seven days. These existing devices attempt to provide total flow andpressure capacity, forcing the pump to operate in inefficient flowregimes.

Centrifugal or radial flow blood pumps are generally somewhat largerthan axial flow devices and provide non-pulsatile flow, but therotational speeds are generally much slower (2,000-10,000 rpm) thanaxial flow blood pumps. While axial flow blood pumps are the smallestVAD, they are higher speed lower pressure rise devices, whilecentrifugal VADs are better suited to take over heart function and toprovide total pressure rise and flow (about 120 mmHg and 5 L/min).Examples are the Gyro C1E3 of Kyocera Corporation (Kyoto, Japan) whichevolved into the NEDO PI-601 pump (animal studies).

Third-generation VADs are those that have replaced the mechanicalbearings of second generation ones with hydrodynamic ormagnetic-suspension bearings. Examples of axial flow VADS are: theINCOR® LVAD of Berlin Heart AG (Berlin, Germany); the MicroVad currentlyunder development at Helmholtz-Institute for Biomedical Engineering(Aachen, Germany); and the MagneVAD I and II of Gold MedicalTechnologies, Inc. (Valhalla, N.Y., USA). Examples of centrifugal flowVADs are: the HVAD of HeartWare Ltd (Sydney, NSW, Australia); theEVAHEART™ of Evaheart Medical USA, Inc. (Pittsburgh, Pa., USA); theVentrAssist LVAD of Ventracor Ltd (Chatswood, NSW, Australia); theCorAide™ LVAD of Arrow International (Reading, Pa., USA); the DuraHeartof Terumo Heart, Inc. (Ann Arbor, Mich., USA); the HeartQuest VAD ofWorldHeart, Inc. (Oakland, Calif., USA); the HeartMate III of ThoratecCorporation (Pleasanton, Calif., USA); and the MiTiHeart™ LVAD of MohawkInnovative Technology, Inc. (Albany, N.Y., USA). All the above devicesrequire major sternotomy or otherwise invasive surgery and CPB.

Other examples of previous devices can be found in the followingpatents, each of which is hereby incorporated by reference: U.S. Pat.No. 4,625,712; U.S. Pat. No. 4,779,614; U.S. Pat. No. 4,846,152; U.S.Pat. No. 5,267,940; U.S. Pat. No. 6,632,169, U.S. Pat. No. 6,866,625;U.S. Pat. No. 7,238,151; U.S. Pat. No. 7,485,104; U.S. Pat. No.8,075,472; U.S. Pat. No. 8,371,997; U.S. Pat. No. 8,545,380; U.S. Pat.No. 8,562,509; U.S. Pat. No. 8,585,572; U.S. Pat. No. 8,597,170; U.S.Pat. No. 8,684,904; U.S. Pat. No. 8,690,749; U.S. Pat. No. 8,727,959;U.S. Pat. No. 8,734,508; U.S. Pat. No. 8,814,933; U.S. Pat. No.8,870,552; U.S. Pat. No. 8,900,115; U.S. Pat. No. 8,961,389; U.S. Pat.No. 9,028,392; U.S. Pat. No. 9,107,992; U.S. Pat. No. 9,138,518; U.S.Pat. No. 9,162,018; U.S. Pat. No. 9,211,368; U.S. Pat. No. 9,295,550;U.S. Pat. No. 9,339,597; U.S. Pat. No. 9,364,593; U.S. Pat. No.9,370,613; U.S. Pat. No. 9,387,285; U.S. Pat. No. 9,474,840; U.S. Pat.No. 9,555,175; U.S. Pat. No. 9,572,915; U.S. Pat. No. 9,579,433; andU.S. Pat. No. 9,597,437.

SUMMARY

It is an object of the invention to provide a device that can beinstalled with less risk to the patient, which reduces disruption tonormal functioning of the heart and/or which minimizes damage to theblood.

According to an aspect of the invention, there is provided a mechanicalcirculatory support, comprising: a body portion defining an internallumen; an inlet port in fluid communication with the lumen; an outletport in fluid communication with the lumen; and a pump for driving fluidflow from the inlet port towards the outlet port, wherein: the inletport is arranged to provide a connection, or is in a state ofconnection, into the aorta of a human body.

This arrangement does not require any connections to be made directly tothe heart and can be installed using minimally invasive surgery, greatlyreducing the risks associated with installation relative to arrangementsthat need to be connected directly to the heart. There is no need toperform a cardiopulmonary bypass for example. The reduced installationrisk makes the device more suitable for treatment of earlier stage CHFthan existing MCS/VAD devices, for example early stage IV CHF. In someembodiments, the device may be suitable for treating stage III or stageIV CHF. The device may be particularly suited to treat late stage IIICHF or early stage IV CHF.

The outlet port may be connected to a downstream position in the aortaso as to be connected in series with the native heart. This type ofconnection is less disruptive to the normal functioning of the heartthan systems which work in parallel with the heart and may help topromote regeneration of the heart muscle. Additionally or alternatively,by allowing the native heart to pump to its best capacity the additionalpumping power required by the support may be reduced.

In an embodiment, the series connection is implemented by connecting thesupport in parallel with a small section of the descending aorta. In analternative embodiment, the descending aorta is interrupted so that allof the blood flow passes through the support.

In other embodiments, the outlet port is connected at other positions inthe vasculature, for example in the ascending aorta. In an embodiment,the support comprises one outlet port in the descending aorta and oneoutlet port in the ascending aorta. In this way, a proportion of theoutflow is provided to the ascending aorta to support coronary flow moredirectly. In an embodiment, the inlet port is connected to one or moreother strategic locations such as the ascending aorta, and the outletport(s) connected as previously described into the descending aorta, theascending aorta, or both. The descending aorta outlet has additionaladvantages for renal, splanchnic, and other organ perfusion withoutaffecting brain flow.

In an embodiment, the pump is a centrifugal pump. The inventors havediscovered that such pumps can provide particularly effective impetus tothe circulating blood. In particular, unnecessary blood shear andfluid-dynamic diffusion (the effect of pressure rise as flow deceleratesalong the device passage) and turbulence can be minimized, which in turnminimizes the imposed shear stress to blood cells, thus minimizing bloodcell lysis (haemolysis) and thrombus formation. The improved pumpingefficiency reduces power requirements, enabling the power supply to bemade smaller and more comfortable to carry. In addition, the pump itselfcan be made more compact. In an alternative embodiment, the pump is amixed flow pump (e.g. a pump having characteristics intermediate betweena centrifugal pump and an axial pump). In a still further embodiment,the pump is a helical pump. In a still further embodiment, the pump isan axial pump.

In an embodiment, the pump is configured to provide a continuous, ratherthan pulsatile flow. The inventors have realised that it is notnecessary for the pump to mimic the pulsatile flow imparted by thenative heart, particularly when installed so as to work in series withthe heart. The pump can thus interact more smoothly with the blood flow,further minimizing damage to the blood. Additionally, the efficiency ofa continuous pump can be optimized further than a pulsatile pump.Acceleration and deceleration of the blood is reduced, which reduces thestresses that need to be applied to the blood as well as the neededpower input to the pump. In alternative embodiments the pump isconfigured to provide a pulsatile flow (synchronous or asynchronous ordifferent fixed phase or variable phase with the heart).

In an embodiment, the support comprises a power receiving member that isconfigured to receive power for driving the pump transcutaneously, forexample by electromagnetic induction. Alternatively or additionally,power can be supplied percutaneously.

According to an aspect of the invention, there is provided a mechanicalcirculatory support, comprising: a pump configured to be installed, orin a state of installation, in a human body and configured to operate inseries with the native heart; and a device for electromagneticallydriving the pump that is configured to be mounted to the body. Thus, asupport is provided that is suitable for “permanent” installation (e.g.so that the patient can leave the hospital with the support installedand operational) and which provides a pumping action that is in series,rather than in parallel, with the native heart.

MCSs which generate full physiological pressure rises (about 120 mmHg),such as VADs in-parallel with the heart, may impart tremendous damage tothe blood (e.g., haemolysis), especially in later stages of CHF. MCSswhich are installed in-series with the heart (i.e. the left ventricle)may exploit the existing pressure rise of the native heart and providean additive pressure rise. Disclosed herein are embodiments of MCSsconfigured for in-series installation in the aorta, particularly thedescending aorta. Installation within the descending aortaadvantageously is conducive to installation via minimally invasivesurgery (e.g., percutaneous installation or thoracoscopy), whichproduces better outcomes (e.g., reduced morbidity) and shorter recoveryperiods for patients, especially those suffering CHF. Additionally,minimally invasive surgical procedures may generally be performed atdistrict hospitals by vascular surgeons, unlike the sternoscopyprocedures that are generally necessary for installation of VADs, whichusually must be performed by cardiothoracic surgeons in critical careunits. Installation within the descending aorta is further advantageousbecause the MCS intercept location is downstream of the cerebral bloodflow, fed by the carotid arteries, reducing the risk of cerebralthromboembolism or stroke. Any blood damaged by an MCS installed in thedescending aorta is pumped to the renal inflow arteries and remainingsystemic and pulmonary perfusion system prior to reaching the cerebralblood flow. MCSs which are installed in the descending aorta must becareful not to establish such a large pressure rise that upstream bloodperfusion to the cerebral blood flow is not suppressed, or stolen, bythe suction of the MCS.

MCSs may be designed with operating conditions specifically configuredfor particular stages of CHF. For instance, a MCS designed for latestage II or early stage III CHF may provide a 20-50 mmHg pressure rise,while a MCS designed for late stage III or early stage IV CHF mayprovide a 40-80 mmHg pressure rise, to better supplant the failingheart. The reduced pressure requirements of MCSs that are installedin-series with the heart may effectively reduce the load on the heart(afterload reduction) by lowering the resistance to blood flow, whichcan advantageously provide the heart increased potential forregeneration of diseased tissue. MCSs with less than full physiologicalpressure rises generally will require less power and will be smaller andlighter weight than MCSs such as VADs which generate larger pressurerises. MCSs installed in series may be configured to maintain thephysiological flow rate of a healthy individual of about 5 L/min. TheMCSs may pump blood at a continuous flow, while the native heart maymaintain pulsatility in total perfusion. In alternative embodiments, theMCS may provide a pulsatile flow. Such pulsatile flow may beestablished, for example, by axially oscillating the impeller within theMCS casing.

Turbomachines operate efficiently over only a very narrow regime ofpressure rise, flow rate and rotational speed specifications, all ofwhich translate into a narrow regime of optimal angles of attack (angleof incoming flow) to turbomachinery airfoils. Therefore, a turbomachineconfigured, for example, to generate a 120 mmHg pressure rise, such as aVAD designed for in-parallel implantation with the left ventricle, willoperate substantially less efficient if instead installed in thedescending aorta and operated at a much lower pressure differential(e.g., 70 mm Hg). For instance, operating a turbomachine below itsconfigured pressure differential will: operate at a much different thanas-designed pressure rise, flow rate, and rotational speed; operate awayfrom the as-designed optimal condition for angles of attack toturbomachine blades; will not work efficiently; and will createunnecessary blood shear, turbulence, stall and losses. These deviationsfrom optimal as-designed operating conditions will increase blood traumaand reduce device efficiency and efficacy for use in this location.

Disclosed herein are embodiments of MCS devices and systems along withmethods of installing and/or using MCS devices to treat CHF. In variousembodiments, the MCS is a centrifugal pump, comprising an impellersuspended in a casing, an inlet introducing blood flow from the nativevasculature to the impeller in an axial direction, and a diffuser withan entrance positioned along the circumference of the impeller and anoutlet returning blood flow to the native vasculature. The impeller maybe magnetically suspended in a contactless manner within the casing androtated using an electromagnetic motor. An external controller implantedwithin the body may provide power to the MCS and control the electricaloperations. The MCS may be powered by internal and/or externalbatteries. The internal batteries may be recharged and/or power may bedelivered from external batteries through transcutaneous or percutaneousenergy transfer systems. In various embodiments, the MCS is specificallysuited for late stage III and/or early stage IV CHF and generatespressures rises between about 40 to about 80 mmHg and maintains a flowrate of approximately 5 L/min.

BRIEF DESCRIPTION OF THE DRAWINGS

Embodiments of the invention will now be described, by way of exampleonly, with reference to the accompanying drawings in which correspondingreference symbols indicate corresponding parts, and in which:

FIG. 1 depicts a mechanical circulatory support connected to a sectionof vasculature and configured to drive fluid flow in parallel with asmall portion of the native blood vessel.

FIG. 2 depicts an alternative configuration for the mechanicalcirculatory support of FIG. 1 in which the support drives blood flowthat is entirely in series with the native blood vessel, bypassing ashort portion of the native blood vessel.

FIG. 3 depicts a mechanical circulatory support comprising multipleoutlet ports and impedance setting members.

FIG. 4 schematically illustrates various installation configurations ofVADs in the vasculature.

FIGS. 5A-5D illustrate an example of an MCS. FIG. 5A illustrates aperspective view of an MCS. FIG. 5B depicts a photograph of an MCSprototype. FIG. 5C illustrates a side cross-sectional view of the MCS.FIG. 5D schematically illustrates a simplified side-cross-section of theMCS 100 along with example dimensions (in mm) of various components andspacing.

FIGS. 6A-6E illustrate an example of an impeller. FIG. 6A illustrates aperspective view of an example of an impeller configured to be used withan MCS. FIG. 6B illustrates a side cross section of the impeller. FIG.6C illustrates a top cross section of the impeller. FIG. 6D illustratesa perspective view of an impeller assembly including a top cap and abottom cap. FIG. 6E illustrates an exploded view of the impellerassembly in FIG. 6D.

FIGS. 7A-7D illustrate perspective views of further examples ofimpellers. FIG. 7A illustrates an example of a shrouded impeller. FIG.7B illustrates another example of a shrouded impeller. FIG. 7Cillustrates an example of an unshrouded impeller. FIG. 7D illustratesanother example of an unshrouded impeller.

FIGS. 8A-8E illustrate examples of an MCS casing. FIG. 8A illustrates anexploded view of an example of an MCS casing. FIG. 8B illustrates abottom view of the casing upper volute shown in FIG. 8A. FIG. 8Cillustrates a perspective view of the casing lower volute shown in FIG.8A. FIG. 8D illustrates a perspective view of another example of an MCScasing. FIG. 8E illustrates an exploded view of an example of an MCSimpeller with inner and outer casings.

FIG. 9 schematically illustrates an example of blood flow through theimpeller and internal casing surface of an MCS.

FIGS. 10A-10D illustrate example components of an MCS magnetic axialsuspension system. FIG. 10A illustrates an example of the relativepositioning of axial-suspension magnets. FIG. 10B illustrates an exampleof an upper axial magnet holder. FIG. 10C illustrates an example of alower axial magnet holder. FIG. 10D schematically illustrates theadjustability of the axial magnet holders relative to the ring magnetspositioned on an MCS impeller.

FIGS. 11A-11E illustrate example components of an MCS magnetic radialsuspension system. FIG. 11A illustrates an example of the relativepositioning of radial suspension magnets and eddy current sensors. FIG.11B illustrates an example of a top radial magnet holder. FIG. 11Cillustrates an example of a bottom radial magnet holder. FIG. 11Dillustrates an example of the upper radial suspension components seatedon an MCS casing lid. FIG. 11E illustrates an example of the lowerradial suspension components seated on an MCS casing lower volute.

FIGS. 12A-12B schematically illustrate two modes of stabilizing animpeller within the casing of an MCS. FIG. 12A illustrates stabilizationusing a passive magnet and hydrodynamic journal bearing force. FIG. 12Billustrates stabilization using passive and active magnets.

FIGS. 13A-13B schematically illustrate the electrical operation of theelectromagnetic stabilization system. FIG. 13A schematically illustratesa block diagram depicting the electrical operation of an electromagneticstabilization system. FIG. 13B schematically illustrates an example of acircuit that may be used according to the flow chart depicted in FIG.13A to operate the electromagnetic stabilization system.

FIGS. 14A-14B illustrate an example of a MCS rotor. FIG. 14A illustratesa top view of the rotor. FIG. 14B illustrates a perspective view of therotor installed within the impeller of an MCS.

FIGS. 15A-15B illustrate an example of a MCS stator. FIG. 15Aillustrates a top view of the stator. FIG. 15B illustrates thepositioning of the stator around an impeller as well as the relativepositioning of the lower axial and radial suspension components.

FIGS. 16A-16F illustrate examples of MCS power systems and operatingparameters. FIG. 16A schematically illustrates an example of atranscutaneous energy transfer system. FIG. 16B schematicallyillustrates an example of a percutaneous energy transfer system. FIG.16C schematically illustrates an example of motor driving circuit. FIG.16D schematically illustrates an example of a battery charging circuit.FIG. 16E schematically illustrates an example of a power conditioningcircuit. FIG. 16F depicts computational results of haemolysissimulations relative to other devices.

FIG. 17 schematically illustrates an example of an MCS installedin-series with a portion of the descending aorta in an angledconfiguration.

FIGS. 18A-18B schematically illustrate examples of an MCS installedin-parallel with a portion of the descending aorta in angledconfigurations. FIG. 18A shows an MCS installed using straight grafts.FIG. 18B shows an MCS installed using two curved grafts.

FIG. 19 schematically illustrates an example of an MCS installedcollinear with a portion of the descending aorta using a question-markshaped outlet graft.

FIG. 20 schematically illustrates an example of a coaxial MCS comprisinga 90 degree flow turn at the inlet installed in-series with a portion ofthe descending aorta.

FIGS. 21A-21D schematically depicts simulated blood flow through variousMCS configurations. FIG. 21A illustrates a MCS installed in an angledconfiguration with approximately 45 degree inlet and outlet anglesrelative to the aorta. FIG. 21B shows a MCS installed in an angledconfiguration with an approximately 65 degree inlet angle and anapproximately 25 degree outlet angle relative to the aorta. FIGS. 21Cand 21D show coaxial MCSs with 25 mm and 15 mm inlet radii,respectively, or MCSs installed in angled configurations with anapproximately 90 degree inlet angle and an approximately collinear (0degree) outlet relative to the aorta.

FIGS. 22A-22C schematically illustrate an example of a collinear MCSwith a wrap-around diffuser and volute passage. FIG. 22A illustrates aside cross-sectional view of the impeller, a portion of the diffuser,and the direction of fluid flow through the diffuser. FIGS. 22B and 22Cillustrate different perspective views of the collinear MCS withwrap-around diffuser and the direction of fluid flow through the MCS.

FIGS. 23A-23E schematically illustrate examples of vanes positionedwithin the inflow or outflow paths of an MCS for altering fluid flow.FIG. 23A schematically illustrates a side cross-section of an example ofan inlet of a device comprising stationary pre-swirl vanes. FIG. 23Bschematically illustrates a side view of the opened circumference ofanother example of an inlet comprising stationary pre-swirl vanes. FIG.23C schematically illustrates a top cross-sectional view of a casingcomprising a splitter vane within the diffuser and volute. FIG. 23Dschematically illustrates a top cross-sectional view of a casingcomprising a splitter vane within the outlet volute. FIG. 23Eschematically illustrates a top cross-sectional view of a casingcomprising diffuser vanes circumferentially positioned around thediffuser.

DETAILED DESCRIPTION

FIG. 1 depicts a section of vasculature 2. In an embodiment, the sectionof vasculature 2 comprises a section of the descending aorta. In anembodiment, the section of the descending aorta is below the diaphragm(arrow 4). In an embodiment, the section of the descending aorta isupstream and/or above the renal arteries and/or splanchnic arteries(arrow 6). Blood flow is shown schematically by arrows 8, 8A and 8B.

A mechanical circulatory support 10 comprises connections into (i.e.through the wall of) the vasculature via inlet port 12 and outlet port14. The inlet port 12 is in fluid communication with a first end 16 of alumen 20 defined by body portion 24 of the support 10. The outlet port14 is in fluid communication with a second end 18 of the lumen 20. Apump 22 is provided within the lumen 20 and configured for driving fluidflow in a direction away from the inlet port 12 and towards the outletport 14.

In an embodiment, the pump 22 is a centrifugal pump. The geometry ofcentrifugal pumps appears at first sight to be less convenient than thatof axial pumps, which are used in some prior art MCS/VAD devices.However, the inventors have recognised that fluid-flow and turbomachineefficiencies gained from using centrifugal impellers, as opposed toaxial impellers, at the selected pressure rise, flow rate, rotationalspeed, and device diameter, as well as from the less aggressiveinteraction between the pump and the blood for a given level of pumpingmore than outweigh any difficulties imposed by the geometry. Levels ofpumping that are required in the context of pumping blood can beprovided with less input power and less damage to the blood. Operationin-series in the described anatomic location results in lower powerlevels than devices designed as VADs configured to provide the full 120mmHg pressure rise, and makes it possible to reduce the dimensions ofthe pump. Reducing damage to blood reduces the risk of adverseside-effects during use.

In an embodiment, the pump 22 is configured to provide a continuousflow, rather than a pulsatile flow (such as that provided by the nativeheart). The resulting pump 22 is simpler and can be optimized moreeasily. The inventors have recognised that it is not necessary to mimicthe pulsatile flow of the heart. This is particularly the case when thesupport 10 is provided in series with the heart because the extent towhich the operation of the support disrupts the normal functioning ofthe heart is reduced in comparison to prior art arrangements that areconnected directly to the heart and arranged to operate in parallel withthe heart.

In the embodiment shown in FIG. 1, the inlet port 12 is configured todivert a portion 8A of the blood flow within the blood vessel into thelumen 20 while allowing the remaining blood flow 8B to continue throughthe native blood vessel 2. The outlet port 14 is configured to allow thereintroduction of the diverted portion 8A of the blood flow back intothe blood vessel 2 further downstream. In this embodiment, the support10 therefore operates in parallel with a short portion 26 of the bloodvessel 2. This approach minimises disruption to the existing vascularsystem and can be installed using minimally invasive surgery. Inaddition, the provision of a region having parallel flow paths increasesthe overall flow capacity of the vascular system, thereby reducing theload on the heart to a degree. The resistance and impedance of segment8B may need to be adjusted to prevent recirculating flow between theoutlet and the inlet of the pump.

In an embodiment, a device is provided for driving the pumpelectrically. In an embodiment, the device is configured to be mountedto the body (e.g. having components that are mounted inside the body,outside the body, or both). The support can thus be installed for longperiods of time (e.g. multiple weeks, months or years). The patient isthus not required to remain within a hospital ward after the support isinstalled. In the embodiment shown in FIG. 1, the device for driving thepump comprises a power receiving member 50, which receives power fordriving the pump. The power receiving member 50 is configured to receivean input of power 52 from a power source located outside of the body(e.g. a battery mounted on the outside of the body) and/or a powersource located inside the body (e.g. a battery mounted inside the body).In an embodiment, the connection between the power source and the powerreceiving member 50 is made wirelessly, for example usingelectromagnetic induction. In an embodiment, the power receiving member50 comprises a coil. Where the wireless connection is made to a powersource outside of the body, the connection may be referred to as atranscutaneous connection. In an embodiment, a wired connection is madebetween a power source located outside the body and the power receivingmember 50. In an embodiment, the wired connection is establishedpercutaneously.

In an embodiment, the support 10 further comprises a datatransmitter/receiver 54 for transmitting/receiving data 56 to/from acontroller 57 outside of the body. In an alternative embodiment, thecontroller 57, or a part of the controller 57, is configured to beinstalled within the body (i.e. under the skin). In an embodiment ofthis type, the controller 57 is sealed in a manner suitable forinstallation within the body and/or comprises a housing made from amaterial that is suitable for being in contact with tissue within thebody for a prolonged period of time (e.g. a biocompatible material). Inan embodiment, the controller 57 comprises a housing made from the samebiocompatible material as a housing for an internal power source (e.g.internal batteries) for powering part or all of the support 10.

In an embodiment, the controller 57 is configured to interact with oneor more sensors for monitoring one or more operating characteristics ofthe pump 22. For example, speed sensors can be used to measure therotational speed of an impeller of the pump 22. In one embodiment three(3) Hall-effect sensors are used to measure impeller rotational speed.Alternatively or additionally, the pressure rise across the impeller ismeasured, for instance with two pressure transducers, one upstream andone downstream of the impeller. In an embodiment, the flow rate ismeasured, or calibrated as a function of other measured parameters. Inan embodiment the set of measurements output from the sensors, or anysubset of the measurements (e.g., impeller rotational speed and pressurerise) are used (for example by the controller 57) to adaptively controlthe rotational velocity of the impeller and therefore also the powerinput to the pump motor in order to achieve the required perfusion. Inother embodiments, other operational characteristics are adaptivelycontrolled in response to one or more sensor measurements.

In one embodiment, performance data, such as impeller rotational speedand/or pressure rise and/or flow rate is/are transmitted to an internalor external unit (e.g. the controller 57 or a part of the controller 57)that is configured to sound an alarm in case of acute conditionsdeveloping, or in case of a system malfunction. In an embodiment, theperformance data is transmitted wirelessly to an external unit thatcollects the data in an application installed in a smartphone or similardevice by the patient's bedside, and for example sends themelectronically to a monitoring station. In an embodiment, the monitoringstation is set up to send an alarm to the patient's guardian orphysician, or to emergency services. Alternatively or additionally, thesystem may be set up to intelligently tune operation of the pump toimprove performance. Further details of the electrical operation of themechanical circulatory support are described elsewhere herein.

FIG. 2 illustrates an alternative embodiment in which the mechanicalcirculatory support 10 is configured to bypass a portion of the bloodvessel 2, rather than operate in parallel with this portion of the bloodvessel 2, as in the embodiment of FIG. 1. The inlet port 12 in thisembodiment diverts all of the flow 8 within the blood vessel 2 into thelumen 20 of the support 10. Similarly, the outlet port 14 is configuredto reintroduce all of the flow 8 back into the native blood vessel 2.Specific examples of mechanical circulatory supports installed eitherin-series and in-parallel with the aorta will be described herein.

In the embodiments described with reference to FIGS. 1 and 2, thesupport 10 has a single inlet port 12 and a single outlet port 14.However, this is not essential. In alternative embodiments, the support10 may comprise two or more inlet ports 12 and/or two or more outletports 14. In an embodiment, the support 10 comprises a single inlet port12 within the descending aorta and two outlet ports 14. In anembodiment, the first outlet port 14 is configured to be connected intothe descending aorta and the second outlet port 14 is configured to beconnected into the ascending aorta. In an embodiment, the support 10 hasa single inlet port 12 connected into the descending aorta and a singleoutlet port 14 connected into the ascending aorta. Providing an outletto the ascending aorta may be useful for example to provide additionalsupport to the brain, or to ‘prime’ the pump. Other configurations arepossible according to clinical need.

Where a multiplicity of outlet ports 14 are provided, flowcharacteristics associated with each of the different outlet ports 14and/or flow paths leading to the outlet ports 14, may be chosen so as tocontrol the distribution of blood flow provided by the pump 22 accordingto clinical need. The flow characteristics may include the flowresistance, flow compliance and/or flow inductance. For example, whereonly a small contribution to the flow is required at a particular outletport 14, the flow resistance associated with that outlet port 14 may bearranged to be relatively high. Conversely, where a relatively high flowoutput from the outlet port 14 is required, the flow resistanceassociated with that outlet port 14 may be arranged to be relativelylow. FIG. 3 illustrates, highly schematically, such a configuration.Here, support 10 comprises a single inlet port 12 and three differentoutlet ports 14A, 14B, 14C. Outlet port 14A is positioned downstream ofthe inlet port 12 in the same section of vasculature 2. The other outletports 14B and 14C are located elsewhere in the vascular system and arenot shown in FIG. 3. Flow characteristic setting members 28A, 28B, 28C,which may be valves for example or sections of tubing of controlleddiameter, are positioned on respective flow paths between the pump 22and each of the three outlet ports 14A, 14B, 14C. By varying the flowcharacteristics using the flow characteristic setting members 28A, 28B,28C, it is possible to define the proportion of the total flow output bythe pump 22 that will be present in the respective flow paths 30A, 30Band 30C.

In an embodiment, the pump is configured to provide a pumping outputthat is equivalent to or greater than the total pumping requirement ofthe body within which the support is installed, so that no additionalpumping from the native heart is required. In an embodiment, the pump22, 34 is configured to provide a pressure of at least 125 mmHg and/orflow rates equivalent to the normal cardiac output of 5 litres perminute. The centrifugal pump approach of the present invention allowssuch pressure and flow rates to be achieved in a compact device withminimum damage to the blood. In another embodiment, the pumping outputis lower than the total pumping requirement of the body. In such anembodiment the pump assists the native heart, which must provide aportion of the total pumping power.

FIG. 4 schematically depicts the differences in installation of variousdevices within the vasculature, including a VAD installed in-parallelwith the left ventricle and outflow connected to the ascending aorta(P1), a VAD installed in-parallel with the left ventricle and outflowconnected to the descending aorta (P2), an MCS installed in-series withthe ascending aorta (S1), and an MCS installed in-series with thedescending aorta (S2), where “MCS” and “VAD” are here used todifferentiate devices installed in-parallel with the left ventricle anddevices installed in-series with the left ventricle, respectively. Asdiscussed elsewhere, each installation configuration may affect theoperating requirements and the installation procedure of the VAD.Installation of a VAD in-parallel with the left ventricle competes forblood flow with the native heart and may essentially take-over pumpingfunction. In-parallel installation may disrupt the natural functioningof the heart and may not allow for full regenerative potential of nativeheart tissue. VADs installed in-parallel may be required to generate thefull physiological pressure rise (about 120 mmHg). VADs installedin-parallel generally need to be installed through highly invasivesurgery (e.g., sternotomy) which generally require performing acardiopulmonary bypass, though there have been recent attempts to modifyinstallation of some VADs to less invasive surgeries, such as describedin Makdisi, G, Wang, I-W., “Minimally invasive is the future of leftventricular assist device Implantation” (2015) Journal of ThoracicDisease 7(9), E283-E288 (incorporated herein by reference). MCSsinstalled in-series add to the pressure rise of the native heart, thusunloading the pressure rise required by the diseased native heart andsupporting its natural function, allowing for regenerative potential ofthe heart. Therefore, because of the lower pressure rise requirement bythe in-series devices, MCSs designed for in-series installation may havelower power requirements. In-series installation of a MCS, particularlywithin the descending aorta, may be performed via minimally invasiveprocedures, without a cardiopulmonary bypass, as the device's flow inletneed not be adjoined directly to the heart. Installation of MCSs withoutlets in the ascending aorta may be used to support cerebral bloodflow. Installation of MCSs with outlets in the descending aorta mayadvantageously avoid risks of blood damage from the MCS causing acerebral thromboembolism or stroke, and they may also increase renalperfusion thus assisting in overcoming cardio-renal syndrome.

FIGS. 5A-5D illustrate an example of an MCS 100. FIG. 5A illustrates aperspective view of the MCS 100. FIG. 5B shows a photograph of aprototype of the MCS 100, demonstrating the approximate size of the MCS100 in a person's hand. FIG. 5C illustrates a side cross-section of theMCS 100. FIG. 5D schematically illustrates a simplified sidecross-section of the MCS 100 along with example dimensions (in mm) ofvarious components and the overall dimensions of the MCS 100. The MCS100 may generally comprise an impeller 200, a casing 300, and magnetholders 402, 404, 406, 408. The casing 300 may include an inlet 102 forreceiving blood flow into the MCS 100, and an outlet 104 for directingexiting blood flow from the MCS 100, both extending from a main body forhousing the impeller 200. The inlet 102 and outlet 104 shown in FIGS.5A-5D are configured particularly for in-vitro testing, and may bemodified accordingly for in-vivo applications (e.g., shortened and/orconfigured for attachment to vascular grafts). The impeller 200 may becontained entirely within the casing 300 and configured to bemagnetically suspended, hydrodynamically suspended, or suspending by acombination of hybrid bearings within the casing 300 such that it doesnot contact the inner surface of the casing 300. The impeller 200 may beconfigured to be electromagnetically rotated within the casing 300 in acontactless manner. The impeller 200 may act as a centrifugal pumpmoving blood received through the inlet 102 from an axial direction andexpelling it centrifugally along the circumference of the impeller 200into the outlet 104. The magnet holders 402, 404, 406, 408 may becoupled to the casing 300 and position magnets and/or electromagnetsaround the casing 300 and impeller 200, which can be used toelectromagnetically suspend and stabilize the impeller 200 within thecasing 300. Other magnets, such as those that drive the rotation of theimpeller 200, may be positioned within the casing 300. As shown in FIG.5B, one or more electrical wires 109 may extend from the MCS 100 (e.g.,they may extend between a controller described elsewhere herein and thecasing 300). The electrical wires may provide power to the device and/ortransmit sensor input to the controller. Each of the operativecomponents of the MCS 100 will be described in further detail elsewhereherein.

FIGS. 6A-6E illustrate examples of the impeller 200 and impellerassembly 201. FIG. 6A illustrates a perspective view of the impeller200. FIG. 6B illustrates a side cross section of the impeller 200. FIG.6C illustrates a top cross section of the impeller 200. FIG. 6Dillustrates a perspective view of the impeller assembly 201, comprisingthe impeller 200, a top cap 207, a bottom cap 209, and other componentsnot visible. FIG. 6E illustrates an exploded view of the impellerassembly 201 depicted in FIG. 6D. The impeller 200 can be configured tobe magnetically suspended within the casing 300 such that the impeller200 is sealed off from the external physiological environment except forblood entering the MCS 100 through the inlet 102. As shown in FIG. 6A,the impeller 200 may comprise a top port 202, a bottom port 204, and amain body 210, each of which may be generally shaped as bodies ofrevolution (e.g., cylindrical). The main body 210 may have a largerdiameter than the top port 202 and/or the bottom port 204. The main body210 may comprise an upper portion 212 (forming an impeller shroud), alower portion 214 (forming an impeller hub), a blade passage chamber 216between the upper portion 212 and lower portion 214, and a plurality ofimpeller blades 218 positioned within the blade passage chamber 216.

As shown in FIG. 6B, the top surface of the upper portion 212 may begenerally open and may extend into an upper chamber 217 configured toreceive a rotor 240, described elsewhere herein, as indicated in FIG.6E. In other embodiments, the lower portion 214 may additionally oralternatively include an open chamber. The outer diameter of the upperchamber 217 may comprise indentations configured to seat and securemagnets of the rotor 240 (FIG. 14B). The upper portion 212 can includean upper channel 203 which may extend from the top surface of the topport 202 to the bottom surface of the upper portion 212 for receivingblood flow into the blade passage chamber 216. The upper channel 203 maycomprise generally circular top and bottom openings. The upper channel203 may be generally cylindrical or frusto-conical in shape, or shapedas a body of revolution to optimize flow patterns at the inlet 102. Theedge between the upper channel 203 and the blade passage chamber 216 maybe generally rounded or curved for directing blood flow in a radiallyoutward direction. A lower channel 205 may extend from the top surfaceof the lower portion 214 to the bottom surface of the bottom port 204.The lower channel 205 may comprise generally circular top and bottomopenings. The lower channel 205 may be generally cylindrical orfrusto-conical in shape. The edge between the lower channel 205 and theblade passage chamber 216 may be slightly rounded to reduce damage tothe blood. The upper channel 203 and/or the lower channel 205 may bealigned generally in the center of the upper and lower portions 212,214. The upper and lower channels 203, 205 may have the same or similardiameters and may be generally aligned with each other in an “axial”direction of the MCS 100, perpendicular to the plane containing theimpeller blades 218 and aligned with the direction blood flow isreceived by the impeller 200.

The bottom surface of the upper portion 212 may form a ceiling to theblade passage chamber 216 and the top surface of the lower portion 214may form a floor to the blade passage chamber 216. The impeller blades218 may extend from the ceiling of the blade passage chamber 216 to thefloor of the blade passage chamber 216 (i.e. between the impeller shroudand the impeller hub). The blades 218 may be integral with the upperportion 212 and the lower portion 214 and may be formed by machining amonolithic piece of material. The impeller 200 shown in FIGS. 6A-6E isan example of a shrouded impeller, as the blades 218 are covered on thetop and bottom by the upper portion 212 and the lower portion 214 suchthat fluid may not flow over or under the blades 218. In otherembodiments, unshrouded impellers may be used as described elsewhereherein. The impeller blades 218 may be generally perpendicular to theceiling and the floor of the blade passage chamber 216 and may form aplane perpendicular to the axial direction of incoming blood flow (theaxial direction of the MCS) in order to facilitate manufacturingconsiderations. In other configurations the impeller blades 218 may bethree-dimensional bodies with lean from the axial direction between thehub and tip (where the blade meets the shroud), in order to optimizeflow parameters. Three-dimensionally shaped blades 218 may be made withadvanced manufacturing techniques such as investment casting orthree-dimensional printing of the biocompatible impeller material. Asshown in FIG. 6C, the blades 218 may each comprise a pressure-side 219and a suction-side 220. The blades 218 may extend in a generally radialor meridional direction from an inner diameter (the leading edge of theblade) to an outer diameter (the trailing edge of the blade). In someembodiments, the blades 218 may be somewhat curved. The pressure-side219 may be convex and the suction-side 220 may be concave, particularlynear the tip of the blade. The inner diameter (leading edge) of theblades 218 may be aligned with the upper channel 203 and/or the lowerchannel 205. The outer diameter (trailing edge) of the blades 218 may bealigned with the outer diameter of the main body 210. In someembodiments, the upper portion 212 and the lower portion 214 may havedifferent diameters and the blades 218 may extend to the larger diameterof the two diameters. The blades 218 may be of a generally uniformthickness as they extend from their leading edge to their trailing edge.In other embodiments, and particularly with advanced manufacturingmethods employed, the blades 218 may be shaped as in modern centrifugalcompressors and radial-inflow turbines of modern turbochargers. The edgeof the blades along the inner diameter (the leading edge) and/or outerdiameter (the trailing edge) may be shaped (e.g., rounded) to match theradius of curvature of the inner circumference or outer circumference,respectively, of the impeller main body 210 (the shroud and/or the hub)to which the blades 218 may be aligned. The shapes of the blades 218along the meridional direction may be shaped with advancedturbomachinery blade-design methods, such as described by T.Korakianitis, I. Hamakhan, M. A. Rezaienia, A. P. S. Wheeler, E. Avitaland J. J. R. Williams, “Design of high-efficiency turbomachinery bladesfor energy conversion devices with the three dimensional prescribedsurface curvature distribution blade design (CIRCLE) method” AppliedEnergy, Vol 89, No. 1, pp.˜215-227, January 2012. (hereby incorporatedby reference). Each of the plurality of blades 218 may be of identicalshape and configuration to the other. The blades 218 may be spaceduniformly around the circumference of the main body 210. The impeller200 may include any number of blades 218 (e.g., three, four, five, six,seven, eight, nine, etc.). Blood flow may be directed from the inlet 102to the blade passage chamber 216 and pumped in a centrifugal directionbetween the blades 218 and out the open circumference portions of theblade passage chamber 216.

FIGS. 6D and 6E illustrate the shrouded impeller assembly 201 inassembled and exploded views, respectively. The top port 202 and bottomport 204 may have the same or similar diameters. The top port 202 and/orthe bottom port 204 may comprise shapes in bodies of revolution. The topport 202 and/or the bottom port 204 may comprise shoulders 211, 213(shown in FIG. 6A) upon which a ring magnet 230 may be seated orpartially seated, as indicated in FIG. 6E. The ring magnets 230,described elsewhere herein, may be configured to slide over the top port202 and/or bottom port 204. In some embodiments, the ring magnets 230may form a tight interference fit with the impeller 200, may be attachedwith advanced joining techniques, or may be fully-inserted into theimpeller material. A rotor 240, described elsewhere herein, may beconfigured to be received within the impeller 200. The impeller assembly201 may further comprise a top cap 207 and/or a bottom cap 209. The topcap 207 and bottom cap 209 may be generally shaped as bodies ofrevolution (e.g., tubular). The caps 207, 209 may comprise flat annularrims extending radially outward at one end configured to be seatedagainst and coupled with the top and bottom surfaces of the main body210, respectively. The caps 207, 209 may have thin annular rimsextending radially inward at the other ends configured to be seated overthe edges of the top port 202 and bottom port 204, respectively. The topcap 207 may be configured to receive the upper port 202 and/or thebottom cap 209 may be configured to receive the bottom port 204 withininner diameters of their bodies. The top cap 207 and/or bottom cap 209may be configured to sit over top of the ring magnets 230 and to sealthem off from the external environment, such as the casing 300. Theradially outward rim of the top cap 207 may be configured to seal theupper chamber 217 and close off the rotor 240 from the externalenvironment, such as the casing 300. In other embodiments, the rotor 240may be positioned in a lower chamber, as described elsewhere herein, oran additional rotor may be positioned in a lower chamber. The top cap202 and/or bottom cap 204 may be coupled to the main body 210 by anysuitable means, including laser welding or a biocompatible adhesive. Insome embodiments, the top cap 207 is contour laser welded to theimpeller 200 and the bottom cap 209 is contour laser welded to theimpeller 200. The impeller assembly may comprise an axial target 221,which may comprise a flat annular right. The axial target 221 may beseated on the bottom surface of the lower portion 214 of the impeller200. The axial target 221 may be fabricated from stainless steel orother suitable materials. The axial target may be magnetic. The impeller200, top cap 207, and bottom cap 209 may comprise a biocompatiblematerial, such as polyether ether keytone (PEEK), for example PEEKOPTIMA, biocompatible titanium, and/or biocompatible titanium coatedwith biocompatible alloys, because they comprise blood-contactingsurfaces.

FIGS. 7A and 7B depict alternative embodiments of impellers 250, 252which exclude top ports and bottom ports. In some implementations, theseimpellers 250, 252 may be subsequently joined to upper and lower portsafter fabrication. As shown in FIGS. 7A and 7B the impellers 250, 252may comprise upper and lower portions 212, 214 of approximately the sameaxial length. In some embodiments, as seen in FIG. 7B, the leading edgesof the blades 218 of impeller 252 may be rounded off. In someembodiments, as seen in FIG. 7B, the leading edges of the blades 218 mayextend inward of the bottom opening of the upper channel 203. Thisconfiguration may allow for easier machining of the leading edges of theblades 218 from the top. Embodiments in which the leading edges of theblades 218 are aligned with the bottom opening of the upper channel 203,as seen in FIG. 6C, may cause less disruption to the incoming bloodflow.

In some embodiments, the impeller may be an unshrouded impeller, asopposed to the shrouded impeller 200 described above. FIG. 7Cillustrates an example of an unshrouded impeller 254 with blades 255that are uncovered on the top and FIG. 7D illustrates an example ofanother unshrouded impeller 256 with blades 257 that are uncovered onthe top. Shrouded impellers have a top (a shroud) and a bottom (a hub)surrounding the impeller blades 218. Unshrouded impellers are uncoveredon one or both sides (top and bottom) of the blades. Fluid may flow overthe tip of the blades 255, 257 in the unshrouded impellers 254, 256illustrated in FIGS. 7C and 7D. Shrouded impellers may have higherefficiencies than unshrouded impellers, due to tip leakage in unshroudedimpellers (i.e. the flow leaks over the rotating blades). Shroudedimpellers introduce more shear to the blood in the region between theshroud and the casing. The MCS may be modified to support an unshroudedimpeller (e.g., with an overhung impeller design). For instance, themotor, comprising the rotor and stator, may be axially positioned aroundthe hub of the unshrouded impeller, rather than around a shroud, and theradial and/or axial stabilization systems (bearings) may also beadjusted appropriately to account for the absence of a shroud. Forinstance, the impeller may be stabilized using the bottom radial andaxial stabilization system components of the impeller along with thestabilization components of the casing, described elsewhere herein.

FIGS. 8A-8E illustrate examples of a casing or components thereof. Thecasing 300 may be configured in shape and dimension to surround theimpeller 200 in such a manner that the impeller 200 may be suspendedwithin the casing 300 and rotated around the axial direction of the MCS100 without any portion of the impeller 200 coming into contact with thecasing 300. The blood contacting surfaces, including casing 300 and theimpeller 200, may comprise one or more biocompatible materials,including but not limited to polyether ether keytone (PEEK), for examplePEEK OPTIMA, biocompatible titanium, and/or biocompatible titaniumcoated with biocompatible alloys. The casing 300 may comprise multiplecomponents which can be assembled around the impeller 200. For example,FIG. 8A illustrates an exploded view of an example of the casing 300.The casing may comprise a lid 312, an upper volute 314, a lower volute316, and an outlet attachment 318. The outlet attachment 318 may beparticularly suitable for in-vitro testing and may be removed ormodified for in-vivo applications, as described elsewhere herein. Thelid 312 may include the inlet 102 or may be joinable to the inlet 102.The outlet attachment 318 can include the outlet 104 and may include acurved section 305 for coupling to the outer circumference of the uppervolute 314 and/or lower volute 316. The components of the casing 300 maybe assembled using screws and/or pins, biocompatible adhesives, or anyother suitable means.

FIG. 8B illustrates a bottom view of the upper volute 314 shown in FIG.8A, and FIG. 8C illustrates a perspective view of the lower volute 316shown in FIG. 8A. The casing 300 can include a diffuser 320. Thediffuser 320 may comprise a passage for receiving blood pumped by theimpeller 200 and may extend into a volute passage in the outlet 104. Thediffuser 320 can be formed directly in the internal surface of thecasing 300, as shown in FIGS. 8A-8C. The diffuser 320 may be formedacross the interface of the upper volute 314 and the lower volute 316.For instance, approximately half the cross-sectional circumference ofthe diffuser 320 may be formed in the upper volute 314 and approximatelyhalf of the circumference may be formed in the lower volute 316. Theupper volute 314 and/or the lower volute 316 may include an indentation315 for receiving a fluid sealing member, similar to an 0-ring, shapedto match the circumference of the diffuser 320. A portion of thediffuser 320 circumference may be open to the internal diameter suchthat blood pumped through the impeller 200 may enter the channel. Inother embodiments, the diffuser 320 may be formed by the addition of acomponent, such as a scroll, along the outer surface of the casing 300,as described elsewhere herein. The diffuser 320 may comprise a partiallycircular cross-section. The diffuser 320 may extend along thecircumferential direction of the MCS 100 to the outlet 104. In someembodiments, the diffuser 320 may simultaneously extend in an axialdirection downward, such that the diffuser 320 begins to spiral. Thediffuser 320 may extend around the entire circumference of the casing300 or only a portion of the circumference. In embodiments in which thediffuser 320 extends around more than a full circumference, the diffuser320 may wrap behind itself closer to the outlet forming an entirelyclosed cross-section, as seen in FIGS. 8A-8C. In some embodiments, thesize of the cross-section of the diffuser 320 may increase as thechannel extends toward the outlet 104. For example, as best seen in FIG.8B, the radial width of the diffuser 320 may continuously increase froman origin point 321 to the outlet 104. The origin point 321 may have avery small thickness such that it forms the beginning of the channelwhich expands in the direction of impeller 200 rotation. In someembodiments, the width of the diffuser 320 may expand along a clockwiseor counter-clockwise direction when viewed from the top. The directionof fluid flow within the diffuser is set by the direction of impellerrotation and blade lean from the radial direction. In some embodiments,the flow-area distribution along diffuser 320 may be chosen to optimizevortex formation in the outlet 104 blood flow. The optimized vortexformation may emulate the weak passage vortex in the healthy nativedescending aorta, as described elsewhere herein.

In various embodiments, the outlet 104 is configured to extendperpendicular to the axial direction of the MCS 100, as shown in FIGS.5A-5D and 8A. The outlet attachment 318 may comprise a volute that formsa continuation of the diffuser 320. The outlet attachment 318 may form asubstantially straight channel. The outlet attachment 318 may provide aconvenient means for attaching an outlet graft which can be anastomosedto the aorta. In some embodiments, the outlet attachment 318 may beexcluded. FIG. 8D illustrates a perspective view of another example of acasing 350 in which the outlet is integral with or contiguous with themain body such that it does not form a cylindrical shaft. In someembodiments, the MCS may comprise multiple layers of casing. FIG. 8Eillustrates an exploded view of another example of a casing 352comprising an inner upper volute 354 and inner lower volute 356, similarto upper volute 314 and lower volute 316, respectively, as well as anouter upper casing 358 and an outer lower casing 360 which areconfigured to surround the inner casing 354, 356 and to interface witheach other along a circumferential seam. In some embodiments, thediffuser 320 may extend into a volute of a scroll ending at the outlet104, as described elsewhere herein. The scroll may further reorientfluid flow, such as by reorienting the fluid flow into a downward axialdirection, such that the MCS may be configured for collinearinstallation within the aorta.

FIG. 9 schematically illustrates in simplified cross-section thesuspended positioning of the impeller 200 within the inner surface ofthe casing 300 and the flow of blood through those components. Thecasing 300 forms a small peripheral space 322 around most portions ofthe impeller 200, excluding the inlet 102 and diffuser 320, each ofwhich forms larger spaces continuous with the primary flow path throughthe impeller 200. The peripheral space 322 allows for contactlessrotation of the impeller 200 by electromagnetic and/or hydrodynamicforces and forms secondary flow paths for blood that fills theperipheral space 322 during operation. The impeller 200 and casing 300form a primary blood flow path, schematically depicted by arrows, fromthe inlet 102 to the diffuser 320 leading to the outlet 104 (not shown).Blood can enter the impeller 200 in an axial direction through the upperchannel 203 and progress through the rotating passages between theimpeller blades 218 which accelerate the blood flow in a tangential andradially outward direction. The blood is forced through the bladepassage chamber 216 (between the blades 218 which are not shown), pastthe outer circumference of the impeller 200, and into the diffuser 320formed in the inner surface of the casing 300. The impeller 200increases the velocity and stagnation pressure of the blood as it passesthrough. The diffuser 320 decelerates the blood flow and increases thestatic pressure. In some implementations, less than half of a generallycircular cross-section defining the diffuser 320 passage may be open tothe internal casing volume containing the impeller 200, as seen in FIG.9. In other embodiments, half or more than half the generally circularcross-section may be open. Although the cross-sections of the diffuser320 on the right and left side of FIG. 9 are shown as equal in size, thecross-sections may be of dissimilar size as the diffuser passage 320 canincrease in cross-sectional area as it extends downstream to the outlet104.

Blood may also flow through secondary blood flow paths, alsoschematically depicted by arrows, formed via the peripheral space 322between the impeller 200 and the casing 300, as shown in FIG. 9. Thesecondary blood flow paths may include an upper secondary blood flowpath and a lower secondary blood flow path. The secondary blood flowpaths may originate in the peripheral space 322 between the bladepassage chamber 216 of the impeller 200 and the casing 300, by flowingupward or downward between the impeller 200 and the casing 300 ratherthan into the diffuser 320. Blood caught in between the impeller 200 andthe casing 300 within the peripheral space 322 provides a hydrodynamicjournal bearing force which helps prevent contact between the impeller200 and casing 300. In an alternative embodiment, the top and bottomflat surfaces of the impeller assembly 201 have spiral grooves, whichbecome part of the secondary flow area in the device gaps, and assistthe hydrodynamic flow through the narrow gaps in order to minimize bloodtrauma within secondary flow paths. Blood may be forced along thesepaths either back to the junction of the inlet 102 and the impeller 200or to the blade passage chamber 216 through the lower channel 205. Thelower volute 316 may include a main stationary shaft 317 (also shown inFIG. 8C) configured to extend from the bottom of the casing 300 into thelower channel 205 of the impeller 200. The main stationary shaft 317 canbe cylindrical or slightly conical in shape, with a correspondingvariation in the shape of the lower channel 205 with which shaft 317forms a hydrodynamic journal bearing. The main stationary shaft 317 maybe configured to reside within the lower channel 205 such that theimpeller 200 can rotate around the shaft 317 in a contactless manner.The upper end of the main stationary shaft 317 may comprise an apex. Theupper end of the main stationary shaft 317 may be shaped to direct flowtoward the circumference of the blade passage chamber 216. The upper endof the main stationary shaft 317 may be flat, conical, conical withconcave surfaces (as shown in FIG. 9), domed, bullet-shaped, rounded, orother suitable shapes. The dimensions of the main stationary shaft 317may be configured to prevent substantial flow in these clearance (gap)areas of the peripheral space 322 rather than along the primary flowpath. The presence of the lower channel 205 allows blood along thesecondary flow path to return to the impeller 200 so that it does notsit stagnant in the residual space around the lower portion 214 of theimpeller 200, thereby enhancing washout of the MCS 100. The axialposition of the impeller may affect the geometry of the flow paths andtherefore the flow rates.

The impeller 200 can be magnetically suspended in the axial directionvia passive (i.e. permanent) magnets positioned within the impeller 200and casing 300. FIGS. 10A-10D illustrate examples of the MCS 100components used to axially suspend the impeller 200. The impellerassembly 201 can include two magnets or two sets of magnets positionedat upper and lower ends of the impeller 200. The casing 300 can includetwo magnets or two sets of magnets positioned at upper and lower ends ofthe casing 300. The impeller 200 can be suspended using the magnets tocreate either approximately equal attractive forces between the impeller200 and the casing 300 at the upper and lower ends of the MCS 100 orapproximately equal repulsive forces between the impeller 200 and thecasing 300 at the upper and lower ends of the MCS 100, accounting forother possible forces such as gravity or accelerations from thepatient's motions. FIG. 10A illustrates an example configuration ofpassive magnets for axial suspension of the MCS 100. The impellerassembly 201 may comprise two ring magnets 230 which can be configuredto be seated around the top port 202 and bottom port 204 (not shown) ofthe impeller 200. The MCS 100 may comprise sets of axial-suspensionmagnets 330 positioned outside the impeller 200. The axial-suspensionmagnets 330 may be positioned within the casing 300, coupled to thecasing 300, and/or positioned between the casing 300 and othercomponents external to the impeller 200, such that the axial-suspensionmagnets 330 remain stationary relative to the housing 300 and physicallyuncoupled from the impeller 200. There may be one or moreaxial-suspension magnets 330 positioned uniformly around the upper andlower circumference of the casing 300. For instance, there may be fouraxial-suspension magnets 330 positioned axially above the upper ringmagnet 230 and four axial-suspension magnets 330 positioned axiallybelow the lower ring magnet 230, as shown in FIG. 10A. In otherembodiments, the axial-suspension magnets 330 may be ring magnetssimilar to ring magnets 230. In an alternative embodiment, the axialsuspension magnets may be positioned slightly further apart in the axialdirection, and by activation via electromagnets coupled to the casing,as described elsewhere herein, be used to axially oscillate the impellerassembly 201 in the casing 300, thus providing pulsatile flow atimpeller outlet.

The upper axial-suspension magnets 330 may be positioned within an upperaxial magnet holder 402, such as that shown in FIG. 10B, and/or thelower axial-suspension magnets 330 may be positioned within a loweraxial magnet holder 404, such as that shown in FIG. 10C. The axialmagnet holders 402, 404 may comprise slots for receiving each of theaxial-suspension magnets 330. The axial-suspension magnets 330 may becoupled to the axial magnet holder 402, 404 via interference fit orother suitable means, such as adhesives, screws, pins, etc. In someembodiments, the upper axial magnet holder 402 may comprise a ring shapeconfigured to fit over the inlet 102, as shown in FIGS. 5A-5D. The upperaxial magnet holder 402 may be secured to the inlet 102 by a frictionfit. The upper axial magnet holder 402 may be slidable along the lengthof the inlet 102 under sufficient force. The lower axial magnet holder404 may be configured as a plate with a central post. The plate may begenerally circular. The post may be generally cylindrical. The post maybe configured to be received within a channel 319 formed generally inthe center of the bottom outer surface of the casing 300 (e.g., thelower volute 316), as depicted in FIGS. 5C and 5D. The length of thechannel 319 may extend into the main stationary shaft 317. The loweraxial magnet holder 404 may be secured to the casing 300 by a frictionfit. The lower axial magnet holder 404 may be translatable within thechannel 319 under sufficient force.

FIG. 10D illustrates the ring magnets 300 coupled to the impeller 200and schematically illustrates the positioning of the upper axial magnetholder 402 and the lower axial magnet holder 404 relative to theimpeller 200. In some embodiments, the ring magnets 230 may be of afirst polarity (e.g., positive or negative). The axial-suspensionmagnets 330 may be of a second polarity, opposite the first polarity,such that the upper ring magnet 230 is pulled axially upward toward theupper set of axial-suspension magnets 330 and the lower ring magnet 230is pulled axially downward toward the lower set of axial-suspensionmagnets 330. In other embodiments, the bottom ring magnet 230 and bottomset of axial-suspension magnets 330 are of a first polarity and theupper ring magnet 230 and the upper set of axial-suspension magnets 330are of a second polarity, such that the upper ring magnet 230 is pushedaxially downward and the lower ring magnet 230 is pushed axially upward.The axial-suspension magnets 330 may be adjustable. For example, asschematically illustrated by the arrows in FIG. 10D, the magnets 330 maybe translatable in an axial direction to modulate the magnetic force andoptimize the axial suspension, as described elsewhere. Positioning theaxial-suspension magnets 330 within the upper axial magnet holder 402and lower axial magnet holder 404 provides for easy axial adjustabilityrelative to the casing 300.

The impeller 200 can be magnetically suspended in the radial directionvia various combinations of passive (i.e. permanent) magnets, active(i.e. electrically activated) magnets or electromagnets (e.g.,conductive coils wrapped around a metal core), and a hydrodynamicjournal bearing effect between the impeller 200 and the internal surfaceof the casing 300. FIGS. 11A-11E illustrate the components that can beused for radial suspension and stabilization. FIG. 11A shows an exampleof the orientation of magnets and sensors used for radial suspension. Apassive radial-suspension magnet 332 may be positioned adjacent to eachimpeller ring magnet 230 (e.g., behind the internal surface of thecasing 300) along the axial direction. The passive radial-suspensionmagnets 332 may be adjustable. For instance, the passive magnets 332 maybe manually translatable in a radial direction such that the passivemagnets 332 may be moved closer to or further from the impeller 200. Insome implementations, the passive magnets 332 may be positioned inmagnet irons comprising an aperture that can be slid or translated alonga rod, pin, or screw in the radial direction. One or more activeradial-suspension magnets 334, described elsewhere herein, may similarlybe positioned adjacent to each impeller ring magnet 230 (e.g., behindthe internal surface of the casing 300). One or more eddy currentsensors 336, described elsewhere herein, may be positioned adjacent toeach impeller ring magnet 230 (e.g., behind the internal surface of thecasing 300). FIG. 11B illustrates an example of a top radial magnetholder 406 and FIG. 11C illustrates an example of a bottom radial magnetholder 408. The radial magnet holders 406, 408 can be used to position(e.g., clamp) the radial-suspension magnets 332, 334 and/or eddy currentsensors 336 adjacent to the casing 300. The radial magnet holders 406,408 may comprise indentations and/or spaces sized to receive orpartially receive the radial suspension components, as shown in FIGS.11B and 11C. FIGS. 11D and 11E illustrate the radial-suspension magnets332, 334 and eddy current sensors 336 seated on the surface of thecasing 300. In some embodiments, the upper and lower outer surfaces ofthe casing 300 are configured to seat all or some of the radialsuspension components. FIG. 11D illustrates the upper radial suspensioncomponents seated on the top of the lid 312. FIG. 11E illustrates thelower radial suspension components seated on the bottom of the lowervolute 316. The casing 300 may comprise identical or similarindentations as the radial magnet holders 406, 408 for partiallyreceiving the radial suspension components, as shown in FIGS. 11D and11E. The components may be sandwiched between the casing 300 and theradial magnet holders 406, 408. The top and bottom radial magnet holders406, 408 may each comprise a ring-like shape configured to be coupledaround generally cylindrical projections extending from the top andbottom of the casing 300, respectively (e.g., the lid 312 and the lowervolute 316). The radial magnet holders 406, 408 may be configured to besecured to the casing 300 by a friction fit or other suitable means.

FIGS. 12A and 12B schematically illustrate two different modes of radialsuspension and stabilization. The impeller 200 may be radially suspendedby the passive radial suspension magnets 332. This can result in radialinstability, according to Earnshaw's theorem, resulting from the axialstiffness. Instability may further result from the magnetic attractionbetween the motor's rotor 240 and stator 340, described elsewhereherein, and from turbulent flow, including vortices, within the MCS 100.The impeller 200 can be further stabilized by journal bearing forcesand/or the active radial suspension magnets 334, as described below.

In some embodiments, as shown in FIG. 12A, a single passiveradial-suspension magnet 332 is used to push the impeller 200 toward theopposite side of the casing 300, creating a large hydrodynamic bearingeffect between the impeller 200 and casing 300. The combined magneticforce between the passive radial-suspension magnet 332 and the impellerring magnet 230 and the journal bearing force may create a radialequilibrium which is highly eccentric, such that the impeller 200rotates around an axis offset from the central longitudinal axis of thecasing 300. This mode of radial suspension advantageously does notconsume additional power because only passive magnets are used andstabilization can be accomplished without additional circuity and/orsensors. In some embodiments, more than one passive radial-suspensionmagnet 332 may be positioned around each impeller ring magnet 230.

In other embodiments, as shown in FIG. 12B, the passiveradial-suspension magnet 332 may be positioned further from the casing300 than the mode depicted in FIG. 12A, such that the impellerequilibrium axis is positioned approximately along the centrallongitudinal axis of the casing 300. Because a less strong journalbearing force is created in this arrangement, the equilibrium point maybe less stable. The active radial-suspension magnets 334 may be used toprevent or inhibit oscillations from the equilibrium point. Eddy currentsensors 336 may be used to monitor the position of the impeller 200, asdepicted in FIG. 12B. The active radial-suspension magnets 334 may beactuated by a control circuit according to input from the eddy currentsensors 336 to stabilize oscillations. The active radial-suspensionmagnets 334 may not act to independently suspend the impeller 200 inorder to limit power consumption. This mode of radial stabilization maybe advantageous because it may result in lower shear stress on theimpeller 200. Lower shear stress may also reduce the amount ofhaemolysis in the pumped blood. Additionally, the active stabilizationallows the MCS 100 to react to dynamic shocks, such as a patient fallingover. In some embodiments, two active radial-suspension magnets 334 maybe positioned around the passive radial suspension magnet 332. Theactive magnets 334 may be positioned on the same side of the impeller200 as the passive magnet 332 and may be symmetrically spaced relativeto the passive magnet 332. Two eddy current sensors 336 may bepositioned on the opposite side of the impeller 200 as the magnets 332,334. Each eddy current sensor 336 may be positioned opposite one of theactive magnets 334. In alternative embodiments, the MCS 100 may rely onone or more other types of bearings to suspend and stabilize theimpeller, including ball bearings, roller bearings, and/or needlebearings.

In some embodiments, the active magnets 334 may be positioned near thering magnets 230 in a position at least slightly axially displaced fromthe ring magnets 230 such that activation of the active magnets 334creates magnetic axial displacement forces between the impeller 200 andthe casing 300. The axial displacement forces may be used to modulatethe axially suspended position of the impeller 200 with respect to thecasing 300. Application of pulsatile phases of current to the activemagnets 334 may be used to oscillate the impeller 200 along an axialdirection and to produce a pulsatile flow. In other embodiments,additional electromagnets distinct from the active magnets 334 may beused to produce the pulsatile flow. In some implementations, theadditional magnets may only be positioned near the upper or lower ringmagnets 230 rather than both.

In some embodiments, the inner axial surface of the casing 300 and/orthe outer axial surface of the impeller 200, or portions thereof, maycomprise circumferential grooves. In some implementations, the groovesmay be spiraled axially. The grooves may have axial gaps between about100 μm and about 1 mm (e.g., 200 μm, 500 μm, 700 μm, etc.). The groovesmay decrease skin friction drag, thereby increasing the efficiency ofthe MCS 100, and may enhance washout flow from the MCS 100. The groovesalso may improve impeller 200 stability by making it easier to axiallysuspend the impeller 200 by adjusting the axial-suspension magnets 330.

FIG. 13A schematically illustrates a block diagram showing an example ofthe circuitry components for operating the magnetic suspension (i.e.maglev) system. FIG. 13B schematically illustrates the circuit asdivided between the four components (blocks 1-4) of the block diagram inFIG. 13A. A conditioning component (block 1) converts and filters theeddy current sensor 336 output into a voltage that can be read by thecontrol circuit. The conditioning component may be a sawtooth generator.The control circuit (block 2) uses the sensor input along with externalinput (the maglev offset) to determine the effort in the correspondingcoils of the active radial-suspension magnets 334. The pulse widthmodulation (PMW) component (block 3) converts the control circuit outputinto a pulse width modulated signal that can be used to drive coilswitching in the active radial suspension magnets 334. The PMW componentmay use comparators. Finally, power MOSFETS (block 4) are driven by thepulse width modulated signal to supply power to the activeradial-suspension magnets 334 configured to stabilize the impeller 200.

The magnetically suspended impeller 200 may be electromagneticallyactuated to rotate around its longitudinal axis within the casing 300via an electromagnetic motor. In some embodiments, the motor may be aradial brushless motor, such as a radial brushless DC motor. The motormay be a radial three-phase brushless DC motor. The motor generallycomprises a stator 340 positioned within the casing 300 and a rotor 240positioned within the impeller assembly 201 and aligned concentricallyinward of the stator 340. FIGS. 14A and 14B depict examples of a rotor240. FIG. 14A shows a perspective view of the rotor 240. FIG. 14B showsa perspective view of the rotor 240 assembled with the impeller 200 inthe impeller assembly 201. The rotor 240 may include passive drivemagnets 242 positioned around a ring 244. The drive magnets 242 may bepositioned on the outer circumference of the ring 244 such that theyextend radially outward from the ring 244. The drive magnets 242 may bepartially embedded within the ring 244. The drive magnets 242 may beuniformly spaced around the circumference of the ring 244. There may beany number of drive magnets 242. In some embodiments, there is a 3:2ratio of stator magnets to drive magnets 242. In some embodiments, theremay be six drive magnets 242. The drive magnets 242 may compriseneodymium (NdFeB). The drive magnets 242 may be generally cubic in shapeand may have dimensions of about 5×5×5 mm. The ring 244 may comprisesteel. The rotor 240 may be configured to be inserted into the impeller200. For example, as shown in FIG. 14B, the rotor 240 may be dimensionedto be inserted into the upper chamber 212 of the upper portion 212 ofthe impeller 200 as described elsewhere herein. The rotor 240 may becoupled to the impeller 200 by any suitable means, including but notlimited to, welding, biocompatible adhesive, or a tight interference fitwith the outer circumference of the top port 202.

FIGS. 15A and 15B depict examples of a stator 340. FIG. 15A shows a topview along the longitudinal axis of a stator 340. FIG. 15B shows aperspective view of the stator 340 positioned around the outercircumference of the impeller 200. The stator 340 may include activemagnets 342 positioned around a ring 344. The ring 344 may comprisesilicon steel. The stator magnets 342 may be positioned on the innercircumference of the ring 344 such that they extend radially inward fromthe ring 344. The stator magnets 342 may be uniformly spaced around thecircumference of the ring 344. There may be any number of stator magnets342. In some embodiments, there is a 3:2 ratio of stator magnets 342 todrive magnets 242. In some embodiments, there may be nine stator magnets342. The stator magnets 342 may comprise metal conductive coils wrappedcircumferentially around projections extending inward from the ring 344.The coils may comprise copper. Electric current provided to theconductive coils may be used to create the electromagnetic forces of theactive magnets. The radially inward end of the projections around whichthe coils are wrapped may comprise circumferentially extending flanges343 which extend towards each other and align with each other to form apartially closed inner diameter configured to sit around an outwardfacing surface of the casing 300 (not shown). Larger gaps may be formedbetween several of the flanges on adjacent projections. The gaps may beconfigured for allowing the positioning of hall effect sensors 346,described elsewhere herein, adjacent to the outer surface of the casing300, as shown in FIG. 15A. In some embodiments, multiple axially-alignedstators 340 (e.g., three stators 340) may be used. The stator 340 may bepositioned within the casing 300. For example, the stator 340 may bepositioned within the upper volute 314.

The motor may be driven by sequentially applying three phases of voltage(positive voltage, zero voltage, and negative voltage) to each statormagnet 342 to induce three phases of current (positive, zero, andnegative) and polarity (positive, non-polar, negative). Pulses ofpositive and negative polarities may travel circumferentially around thestator ring 344 to continuously drive the rotor 240 through magneticinteraction with the drive magnets 242. A controller, which may beexternal to the MCS 100, may be used to time the charging of each statormagnet 342 so as to induce continual rotation of the rotor 240. One ormore bipolar hall effect sensors 346 (e.g., three sensors) positionedwithin the casing 300 may be used to detect the positioning of the rotor240 with respect to the stator 340 by detecting the proximity of a drivemagnet 242. The controller may monitor the output of the one or morehall effect sensors 346 and use the positioning location to modulate theactivation of the stator magnets 342. In some embodiments, the halleffect sensors may be Honeywell part number SS411A sensors.

The electrical systems of the MCS 100 may control the motor and magneticsuspension systems, as well as power conditioning and battery charging.The electrical systems, or a portion of the electrical systems, may beexternal to the MCS 100. The electrical systems may be powered by aninternal rechargeable battery, such as a chemical battery (e.g., lithiumion) or the battery may be used as a backup power source. The internalbattery (or batteries) may be implanted within the body at a positionseparated from the MCS 100 device. For example, the internal batteriesmay be contained in a separate controller device implanted in the body,similar to the manner in which a pacemaker is implanted within a body.The controller may also contain the other electrical systems. In someembodiments, the battery may be charged transcutaneously, via inductivepower transfer through the skin. In some embodiments, the MCS 100 isprimarily powered by an external battery (e.g., a 16.8 V battery), butmay have an internal battery for backup. Power from the external batterymay also be transferred transcutaneously through the skin. FIG. 16A,schematically depicts the components of an example of a transcutaneousenergy transmission system (TETS), including various componentefficiencies (η). An external battery charger may receive line ACvoltage (e.g., 110-240 VAC) and convert it to DC voltage to chargeexternal batteries (e.g., lithium ion batteries). A DC-DC converter maybe used to stabilize the DC voltage provided by the external batteries(e.g., while they discharge). A DC to high frequency (HF) converter mayconvert the DC voltage into a high frequency (e.g., 250 kHz) AC voltagefor transcutaneously charging a secondary coil beneath the skin from anexternal primary coil (e.g., spaced 20 mm apart). Higher frequencies maybe required to transfer energy between coils spaced further apart. Thecoils may be made of Litz wire. An HF to DC converter may be used toconvert the energy back to DC within the body. An internal DC-DCconverter may be used to stabilize the DC voltage supplied to thecontroller. The controller may be electrically connected to the MCS 100(denoted as “TC”) via suitable wiring, including input and outputcapabilities. The controller may include intelligent functioningmechanisms, including constant monitoring of power consumption, impellerrpm, blood pressure, and other performance parameters. Information maybe wireless transmitted to and/or from the controller, such as to apatient, physician, or hospital.

The controller may also include internal rechargeable batteries. Theinternal batteries may serve as temporary backup for when the TETS isdisconnected. The internal batteries may be charged from the output ofthe HF to DC converter. An undercurrent transducer may be used to sensecurrent from the external batteries and switch between power supplieddirectly from the HF to DC converter to power supplied from the internalbatteries, if the current is below a predetermined threshold. Largerbatteries may provide longer independent operation times. Charging thebatteries at lower currents (e.g., 0.2 A) may advantageously limit thetemperature rise of the devices, although longer charging times may beneeded. In some embodiments, the battery may be charged percutaneously.FIG. 16B, schematically depicts the components of an example of apercutaneous energy transmission system (PETS), including variouscomponent efficiencies (η). The MCS 100 may include any suitable meansfor minimizing the electromagnetic interference from other sources,including but not limited to, optimizing the voltage and current for aconstant power, modifying the frequency of the signals, and usingfilters, shields, and/or snubber circuits.

The controller may contain electronic circuitry for operating the MCS100. In some embodiments, the motor can be driven using an L6235 driverchip (ST Microelectronics). FIG. 16C schematically illustrates the L6235driver chip circuit. This circuit can be used to power the hall effectsensors, monitor their output, and drive the three phases accordingly.FIG. 16D schematically illustrates a battery charging circuit. Thebattery charging circuit may use an MSP430 microcontroller to monitorbattery voltage and/or current into the battery via a ZXCT1041 currentmonitor. The microcontroller may stop charging to prevent overchargingif the battery is fully charged and the current into the battery isbelow 0.02 C. Charging may resume when the battery voltage drops below apredetermined threshold. Power into the battery may be controlled by anMMBTA bipolar junction transistor and a BSP250 MOSFET. A variety ofcharging algorithms may be programmed into the microcontroller. FIG. 16Eschematically illustrates a power conditioning circuit. The powerconditioning circuit can be used to create lower voltage levels from thebattery (e.g., a 16.8 V battery) as described elsewhere herein. Runningsome circuits at lower voltages may reduce the power consumption of theMCS 100. Adjustable DC-DC current regulators may be used to ensureefficient conversion. In some implementations, the control electronics,digital filtering, and maglev actuators may be powered at 3.5 V, 5 V,and 6.5 V respectively. In some implementations, the controlelectronics, digital filtering, and maglev actuators may be powered at3.5 V, 3.5 V, and battery power (e.g., 16.8 V) respectively, which mayprovide lower cost, complexity, and power consumption. Electrical powermay be provided from the controller to the MCS 100 via electrical wires109, illustrated in FIG. 5B. There may be multiple wires extendingbetween the controller and the MCS 100. For instance, there may be awire providing power to the radial suspension electromagnets, a wireproviding power to the electromagnets of the motor, a wire receivinginput from the eddy current sensors, a wire receiving input from thehall effect sensor, etc. Power and data may be transferred between thecontroller and the MCS according to any suitable means known in the art.

The MCS 100 may be optimized for performing in-series in a patient withlate stage III and/or early stage IV CHF. The MCS 100 may be optimizedto provide maximum power efficiency, minimize occupying space, and/orreduce device weight. Optimizing power efficiency may reduce batteryweight and/or maximize untethered time during which the device may beoperated via battery power. The device may be configured to optimizestability of the rotating impeller 200 to prevent damage to the deviceand/or blood trauma. Losses in motor efficiency may be electrical,magnetic, and/or mechanical. Electrical efficiency losses may, forexample, include winding resistance (i.e. copper loss), especially inlow speed applications. Magnetic efficiency losses may includehysteresis, eddy current losses, and/or excess eddy current. Mechanicallosses may include windage, ventilation, and/or bearing friction. Insome embodiments, the efficiency is at least 15%. In some embodiments,the efficiency is at least 20%. In some embodiments, the powerconsumption may be about 10 W or less. Efficiency may generally beincreased by using a smaller impeller with reduced skin friction toimprove hydraulic efficiency. Efficiency may generally be increasedallowing more space for coils and/or reducing the stator-rotor gap toimprove electromechanical efficiency at the operating condition.Stability may generally be improved by increasing the stator-rotor gap.

The operating design may be configured to minimize damage to the bloodso that haemolysis is low. Haemolysis is the result of blood traumaimparted by high shear and by time of exposure (or length of flowpassage) in high-shear flow conditions. For a set flow rate (e.g., 5L/min) and to a first approximation, increasing the pressure requireslarger power inputs to the flow and therefore results in larger lossesby friction. Accordingly, the blood trauma imparted by a VAD or MCSincreases as the pressure rises. Therefore, as the MCS 100 is designedto provide 40-80 mmHg, it will result in lower haemolysis than anotherMCS or VAD delivering 5 L/min at much higher pressure rises (e.g.,120-140 mmHg). FIG. 16F depicts the Normalised Index of Haemolysis (NIH,g/100 L) of computation simulations on the MCS 100 (depicted asTURBOCARDIA V5) as well as a prior version having an impeller comprisinglarger upper and lower portions 214, 216 amongst other designdifferences (depicted as TURBOCARDIA V4) and other VADs known in the art(the HVAD and Heartmate II). In some embodiments, as demonstrated inFIG. 16F the computed haemolysis of MCS 100 may be around 0.6 g/100 L.In other embodiments, the computed haemolysis may be less than 0.6 g/100L.

The MCS 100 may be configured for installation within a portion of thedescending aorta. The MCS 100 may be configured to provide approximatelya 40-80 mmHg pressure rise (e.g., about 70 mmHg) at a continuous flowrate of about 5 L/min. The MCS 100 may be configured to operate therotor 240 at approximately 2600 rpm. In some embodiments, the device mayweigh about 150 g. The displacement volume may be about 70 cm³.Referring back to FIG. 5D, example dimensions (in mm) of various MCS 100components and the overall dimensions of the MCS 100 are depicted (theillustrated dimensions may not be drawn to scale). The outer diameter ofthe MCS 100 (around the casing 300) may be between about 30 mm and about100 mm, between about 40 mm and about 70, between about 50 mm and about60 mm, and ranges there between (e.g., about 57 mm). The axial length ofthe casing 300 may be between about 20 mm and about 60 mm, between about30 mm and about 50 mm, between about 35 mm and about 45 mm, and rangesthere between (e.g., about 40 mm), excluding the length of the inlet102. The impeller 200 may have a maximal radial diameter between about10 mm and about 60 mm, between about 20 mm and about 50 mm, betweenabout 25 mm and about 40 mm (e.g., 30 mm). The diameter of the upperchannel 203 may be between about 3 mm and about 25 mm, between about 5mm and about 20 mm, between about 8 mm and about 12 mm, and ranges therebetween (e.g., about 10 mm). In some embodiments, as shown in FIGS. 5C,5D, and 6B the diameter of the upper channel 203 may decrease from theinlet 102 to the blade passage chamber 216. For example, the diameter ofthe upper channel 203 may linearly decrease from about 12 mm to about 8mm. In other embodiments, the upper channel may have a constant diameteror a diameter than decreases in a non-linear manner. The diameter of thelower channel 205 may be between about 3 mm and about 30 mm, betweenabout 5 mm and about 20 mm, between about 8 mm and about 12 mm, andranges there between (e.g., 10 mm). The diameter of the lower channel205 may be constant as shown in FIGS. 5C, 5D, and 6B. In otherembodiments, the diameter may increase in a linear or non-linear mannerfrom the blade passage chamber 216 to the bottom of the impeller 200.The height of the blade passage chamber 216 may be between about 2 mmand about 30 mm, between about 3 mm and about 10 mm, and ranges therebetween (e.g., 5.5 mm). The height of the diffuser 320 may be betweenabout 2 mm and about 30 mm, between about 3 mm and about 10 mm, andranges there between (e.g., 7 mm). In some embodiments, as describedelsewhere herein, the height and/or depth of the diffuser 320 may varydepending on the circumferential position. The gaps between the impeller200 and the casing 300 in the peripheral space 322 may be between about100 μm and 1 mm (e.g., 700 μm). The width of the peripheral space 322may be the same or may vary around different portions of the impeller200 and casing 300. The precise width of the peripheral space 322 maydepend on the operation of the MCS 100, including the axial and radialsuspension, as described elsewhere herein. The inlet 102 may have aninner diameter of about 9 mm. The inner diameter of the inlet 102 may bethe same or less than the diameter of the upper channel 203 where theinlet 102 and upper channel 203 meet. The outlet 104 (not shown) mayhave an inner diameter of about 11 mm. In alternative embodiments, theMCS may be configured for installation in the ascending aorta. The MCSconfigured for installation in the ascending aorta may comprise a secondoutlet which could be configured to send about 5% of the blood flow tothe coronary arteries and the remainder of the blood flow downstream.

The MCS 100 can be installed within the vasculature 2 in variousconfigurations. In various embodiments, the MCS 100 comprises an inlet102 and an outlet 104, which may be arranged generally perpendicular toeach other as described elsewhere herein. The outlet 104 may bepositioned at the end of a diffuser for altering and/or reorienting thefluid outflow. The MCS 100 can be installed into the vasculature usingvascular grafts comprising standard biocompatible graft material (e.g.,polytetrafluorethylene, polyethylene terephthalate, etc.). In someimplementations, patient allografts may be used. The grafts may beconnected to the inlet 102 and outlet 104 of the MCS 100 in any suitablemanner which creates a fluid tight seal. The grafts may be sutured intothe native vasculature.

In some embodiments, the MCS 100 is installed at an angle relative tothe axis of the aorta. For example, FIG. 17 schematically depicts anexample of an MCS 100 installed in-series with the descending aorta, inwhich the inlet 102 and the outlet 104 of the MCS 100 are anastomosed tothe aorta by an inlet graft 106 and an outlet graft 108. The grafts 106,108 may extend from the axis of the aorta at an angle selected from awide array of angles generally between 0 degrees and 90 degrees. Forembodiments of the MCS 100 in which the inlet 102 is substantiallyperpendicular to the outlet 104 (i.e. 90 degrees), the sum of the angleof the inlet 102 relative to the aorta and the angle of the outlet 104relative to the aorta is approximately 90 degrees, when the MCS 100 isinstalled within a generally straight portion of the aorta. For example,as shown in FIG. 17, the inlet 102 and outlet 104 of the MCS 100 areeach arranged approximately 45 degrees relative to the descending aorta.The installation of the MCS 100 within the aorta, particularly at anangle, may somewhat displace or alter the orientation of the upstreamand/or downstream portion of the aorta to which the MCS 100 isanastomosed.

In some embodiments in which neither the inlet 102 nor the outlet 104 ofthe MCS 100 is configured to be collinear with the aorta (the MCS 100 islaterally displaced from the aorta), the MCS 100 may be connectedin-parallel with the aorta. In embodiments where the MCS 100 isconnected in-parallel, the inlet and outlet grafts 106, 108 may beanastomosed with the native vasculature in a branched fashion. In somein-parallel embodiments, the native aorta may be occluded between theinlet graft 106 and the outlet graft 108, effectively making the MCS 100in-series with the aorta. In some in-parallel embodiments, a one-wayvalve (e.g., a one-way artificial heart valve) may be installed in thenative aorta between the inlet graft 106 and the outlet graft 108,permitting blood flow only in the downstream direction. Mechanicallypreventing upstream blood flow within the native aorta mayadvantageously prevent recirculation of blood along a path ofleast-resistance up the native aorta and back through the MCS 100 wheninstalled in-parallel, which may excessively damage the blood and/ordisrupt downstream blood flow.

FIGS. 18A and 18B schematically depict an example of an MCS 100installed in-parallel with the descending aorta. FIG. 18A shows the MCS100 installed at approximately a 60 degree angle between the inlet 102and aorta and approximately a 30 degree angle between the outlet 104 andaorta. FIG. 18B shows the MCS 100 installed at approximately a 90 degreeangle between the inlet 102 and the aorta. The outlet 104 is parallel tothe bottom portion of the aorta (i.e. 0 degrees) and connected via acurved outlet graft 108. In the example illustrated in FIG. 18B, theinlet and outlet grafts 106, 108 are substantially curved. Using curvedgrafts may allow the installation of the MCS 100 in the vasculature atsharper angles and/or may minimize the amount of space occupied by thegrafts 106, 108 and the MCS 100. The curvature of the grafts may alsoeffect vortex formation as described elsewhere herein. The grafts 106,108 may be substantially rigid to support the MCS 100 within thevasculature. Grafts of various shapes or flexibility may be employeddepending on the amount of curvature desired. Embodiments which use moremoderate angles (e.g., 45 degrees) can be advantageous in that theirinstallation can be accomplished using relatively short and/orrelatively straight grafts 106, 108, which may minimize the totalinstallation space of the MCS 100. Use of straight grafts 106, 108 mayimpart less turbulence on the blood flow than use of more curved grafts106, 108.

In some embodiments, the outlet 104 of the MCS 100 is connected to asubstantially curved graft 108 to return blood to the downstream portionof the aorta. The curved outlet graft 108 may extend from the outlet 104of the MCS 100 in a direction substantially perpendicular to the inlet102 and curve toward the downstream portion of the aorta until the graft108 is substantially collinear with the aorta at which point the graftand downstream portion can be anastomosed. FIG. 19, schematicallydepicts an example of a MCS 100 installed in-series with the descendingaorta, in which the inlet 102 is anastomosed to the upper portion of thedescending aorta in a collinear manner or at a relatively small angle(e.g., 0-10 degrees) and the outlet 104 is anastomosed to the lowerportion of the descending aorta via a generally “question mark” shapedoutlet graft 108. This configuration may be advantageous in that itallows installation of the MCS 100 with both the inlet and outlet grafts106, 108 anastomosed to the native vasculature in a generally collinearfashion. Collinear installation of the MCS 100 may minimize the amountof manipulation required in the native aorta to accommodate the MCS 100.Use of an outlet graft 108 with a large radius of curvature may minimizethe amount of turbulence imparted to the blood flow through the MCS 100.

In some embodiments, an MCS 110 may be installed within the aorta in aco-axial configuration, in which the inlet 112 and outlet 114 are notperpendicular but are coaxial, such that they inlet 112 and outlet 114are parallel to a common axis, generally aligned with a longitudinalaxis of the native aorta. FIG. 20 schematically depicts an example of acoaxial MCS 110 installed within the descending aorta. The inlet 112includes a 90 degree bend, allowing the inlet graft 106 to remaincollinear with the upper portion of the descending aorta. Blood flowenters the coaxial MCS 110 impeller from the 90 degree bend “sideways”with respect to a standing patient. The diffuser sends the blood flowvertically downward with respect to a standing patient. Thisconfiguration can result in minimal losses in pump efficiency at theinflow graft 106 as the pressure at that point is relatively lowrelative to other configurations. The remaining features of the MCS 110may be the same as that of MCSs 100 installed in angled configurations.The coaxial configuration may result in the formation of a vortex at theMCS outlet 114. In embodiments comprising a sharp 90 degree bend in theinlet 112, the MCS 110 can be installed with relatively short grafts106, 108 and with minimal installation space. The coaxial MCS 110 may beespecially conducive to installation by minimally invasive surgery. Insome embodiments, the downstream portion of the severed aorta may beslightly displaced upon installation, such as by 3-10 cm, for example.In other embodiments, the outlet 114 may bend to wrap partially aroundthe body of the MCS 110 such that the inlet 114 and outlet 116 arecollinear.

FIGS. 21A-21D schematically depict simulated fluid flow through MCSdevices installed in-series with the aorta in various configurations.FIG. 21A shows a MCS 100 installed in an angled configuration withapproximately 45 degree angles between the inlet 102 and aorta and theoutlet 104 and aorta. FIG. 21B shows a MCS 100 installed in an angledconfiguration with an approximately 65 degree inlet 102 angle and anapproximately 25 degree outlet 104 angle relative to the aorta. FIGS.21C and 21D show MCSs 100 installed in angled configurations with anapproximately 90 degree inlet 102 angles and approximately collinear (0degree) outlets 104 relative to the aorta. The simulations depicted inFIGS. 21C and 21D may be used to approximate the fluid flow through acoaxial MCS 110 comprising a 90 degree bend in the MCS inlet 112. Theexample shown in FIG. 21C has a 25 mm radius at the inlet 112 and theexample shown in FIG. 21D has a 15 mm radius at the inlet 112. Thecoaxial MCSs 110 shown in FIGS. 21C and 21D show no discernible vorticesin the outflow. The angled MCSs 100 shown in FIGS. 21A and 21B showdiscernible vortex formation in the outflow of each. The simulationresults suggest that bending in the outlet may create more fluidvortices than does bending in the inlet. The relatively low pressure atthe inlet 102 and the relatively high pressure at the outlet 104, of theangled MCS devices 100, may stimulate vortex formation. The size of thediffuser at the outlet may also effect vortex formation.

Vortex formation in the outflow of the MCS 100, 110 may be beneficial.For instance, vortex flow may enhance the perfusion of side arteriesbranching from the aorta and/or may enhance washout in the descendingaorta. Using the MCS to recreate physiological flow conditions mayreduce the risk of thrombosis or other pathological conditions. Studieshave shown the identification of right-handed helix formation throughthe ascending aorta and aortic arch into the descending aorta duringsystolic outflow in healthy individuals. See Markl, M. et al. (July2004). Time-Resolved 3-Dimensional Velocity Mapping in the ThoracicAorta: Visualization of 3-Directional Blood Flow Patterns in HealthyVolunteers and Patients, Journal of Computer Assisted Tomography, 28(4),459-468 (incorporated herein by reference). In some embodiments, the MCSand/or the installation of the device may be configured to optimizevortex formation (e.g., to form a right-handed helix) in the outflow ofthe device. For example, the direction of impeller rotation, orientationof the diffuser, inflow angle, outflow angle, inlet diameter, and/oroutlet diameter may be selected to emulate optimal physiologicalconditions, including a weak vortex. Depending on the geometry of theMCS, these parameters may be used to either increase or decrease theamount of vortex formation to mimic that of the native aorta. Prior MCSdevices have aimed to eliminate any vortex formation altogether.

In some embodiments, the MCS is collinear with both the upper portionand the lower portion of the aorta, so that there is no axial or angulardisplacement in the inflow or outflow. FIGS. 22A-22C illustrate anexample of a collinear MCS 120. FIG. 22A schematically illustrate across section of an example of a collinear MCS 120, including animpeller 126 and diffuser 128. FIGS. 22B and 22C illustrate perspectiveviews of an example of a collinear MCS 120. The inlet 122 of the MCS 120may be grafted directly in-line with the upper portion of the descendingaorta. In some variations, the inlet 122 may include pre-swirlstationary vanes (not shown) above the impeller 126, described elsewhereherein. Blood may be pushed by the impeller 126 in a radially outwarddirection into the diffuser 128. The diffuser 128 may reorient theoutflow from a radial direction, aligned 90 degrees relative to theinflow, to an axial direction, aligned collinear with the inflow andwith the lower portion of the descending aorta. The diffuser scroll 129may wrap-around the casing of the MCS 120. The diffuser scroll 129 mayextend inward toward the longitudinal axis of the MCS 120 once itextends below the bottom of the MCS 120 casing. The diffuser scroll 129may extend in a spiral/helical fashion. In some implementations, thediffuser scroll 129 may progressively turn toward the axial direction asit wraps around the casing. The diffuser scroll 129 may gradually shiftflow from a circumferential to an axial direction or may turn to theaxial direction primarily near the outlet 124. The wrap-around diffuser128 sends flow vertically downward and may terminate in a funnel-likeshape at the outlet 124 with an expanding diameter. The diameter of thediffuser scroll 129 may increase as it extends from the impeller 126toward the outlet 124. As seen in the cross-section of FIG. 22A, thecross-section of the diffuser scroll 129 may be smaller on one side ofthe MCS 120 (e.g., the right side of the figure) than the other side(e.g., the left side of the figure). Blood may travel through thediffuser scroll 129 along the direction of the diffuser's increasingsize. The helical direction of blood flow through the diffuser 128 isschematically illustrated by the continuous arrow in FIG. 22A. Theincreasing diameter of the diffuser 128 may promote vortex formation inthe outflow.

The diffuser 128 may perform only a partial revolution around the axisof the MCS 120, a single revolution, multiple revolutions, or any degreeof revolutions there between. For example, the diffuser 128 may make ahalf turn, a three-quarter turn, a whole turn, one and a half turns, twoturns, two and a half turns, three turns, etc., before terminating atthe outlet 124. The azimuthal turning in the scroll 129 from point 321of the diffuser 320 to the end of the turning in the scroll 129 could beany angle or could be at a varying angle. The diffuser 128 may make asharp bend in the axial direction just before reaching the outlet 124.The wrap-around design may be useful for inducing vortex formation inthe outflow of the MCS 120. The design parameters of the diffuser 128may be altered to optimize helix formation. These may include thediameter of the diffuser 128, the change in the diameter of the diffuser128, the number of revolutions made by the diffuser 128, the pitch ofthe turns, and the sharpness in the bend toward the axial direction,particularly toward the outlet. The configuration of the collinear MCS120 may be relatively compact. The wrap-around diffuser 128 may minimizethe overall diameter of the MCS 120. The collinear configuration mayreduce the length of inlet and/or outlet grafts 106, 108, thus reducingthe overall axial length of the MCS 120. The generally small size of thecollinear MCS 120 may make it particularly conducive for installationvia minimally invasive surgery.

The MCS 100 (and other MCSs disclosed herein) may employ stationaryvanes to further alter the inflow and/or outflow of blood through thedevice. In some embodiments, the MCS 100 may include stationarypre-swirl vanes 323 (also known as inlet guide vanes). FIG. 23Aschematically depicts a side view of an inlet 102 comprising stationarypre-swirl vanes 323. FIG. 23B schematically depicts an opened/flattenedcircumferential portion of inlet 102 comprising stationary pre-swirlvanes 323. One or more of these vanes 323 may extend from the innercircumference of the inlet 102 into the axial flow path of theintroduced blood. The vanes 323 may be substantially flat. In otherembodiments, the vanes 323 may have a curved surface. The vanes 323 maycurve the blood flow in the direction of impeller rotation. In someimplementations, the curves may curve the flow in the direction of thenative aortic passage vortex. As shown in FIG. 23A, the vanes 323 maydecrease in width as they extend from the inner diameter of the inlet102 toward the longitudinal axis of the inlet 102. In some embodiments,the vanes may extend to the longitudinal axis. The decreasing width mayallow the accommodation of adjacent vanes 323 around the circumferenceof the inlet 102. As shown in FIG. 23B, the vanes 323 may be angled withrespect to the circumference of the inlet such that they extendpartially in a circumferential direction and partially in an axialdirection. The vanes 323 may all be identical in shape or they may varyin shape. The vanes 323 may all extend at the same angle relative thecircumference and longitudinal axis or they may extend at differentangles. In some implementations, as shown in FIG. 23A, the vanes 323 maybe configured such they cumulatively occupy the entire cross section ofthe inlet 102, but because they are angled blood may flow between thevanes 323. In some embodiments, the vanes 323 may partially overlap eachother in the axial direction. In some embodiments, the vanes 323 do notoccupy the entire cross section of the inlet 102, such that blood couldpotentially flow in a purely axial direction between the vanes 323. Thevanes 323 may pre-swirl the blood entering the MCS 100 prior to reachingthe impeller 200. The vanes 323 may improve fluid dynamics of blood flowthrough the MCS 100 (add a rotational velocity to the blood flow) at thecost of increased friction with the blood. The improved fluid dynamicsmay be used to adjust the flow rate and/or improve the efficiency of theturbomachine. For example, the vanes 323 may allow increased rotationalspeed with reduced motor power. In some embodiments, there may bemultiple rows of pre-swirl vanes 323 along the axial direction. In someembodiments, the vanes 323 may not all be positioned at the same axialposition but may be axially spaced from each other (e.g., in a helicalformation). In some embodiments including pre-swirl vanes 323, pre-swirlvanes 323 may be directly incorporated into the upper channel 203 of theimpeller in addition to or alternatively to the inlet 102. In someembodiments, the vanes 323 may be incorporated into the outlet 104 inaddition to or alternatively to the inlet 102.

In some embodiments, the MCS 100 may include a vaned diffuser 320(and/or a vaned volute extending at the terminal end of the diffuser320). The vaned diffuser 320 may be used to optimize fluid dynamics,such as vortex formation, in the outflow of the device. FIG. 23Cschematically illustrates an example of a top cross-section of a casing300 comprising a diffuser 320 with a single splitter vane 324 whichcreates a split double volute at the outlet 104, comprising two parallelfluid passages. One or more splitter vanes 324 may be used to even outflow distribution, particularly between the inner side of the volute(left side of FIG. 23C) and the outer side of the volute (right side ofFIG. 23C). FIG. 23D, schematically illustrates a variation of the splitdiffuser shown in FIG. 23C, in which the diffuser vane 324 only extendspartially or not at all into the circumferential diffuser 320 passage(the portion of the passage prior to the straight volute passage). Insome embodiments, the splitter vane(s) 324 is not a wall aligned purelywith the axial direction of the device. The splitter vane(s) 324 mayrotate relative to the cross-sectional circumference of the passage asit extends along the diffuser and/or volute. The use of a rotatingsplitter vane(s) 324 may add rotational velocity to the blood outflowand may be used to help emulate the naturally occurring vortex formationin the healthy aorta. FIG. 23E schematically illustrates an example of acasing 300 with a vaned diffuser comprising a plurality of diffuservanes 325 surrounding the inner circumference of the diffuser 320. Thediffuser vanes 325 may be slightly curved in a direction configured toorient the blood toward the outlet 104. The diffuser vanes may beuniformly spaced around the circumference of the diffuser 320. In someembodiments, not all portions of the circumference of the diffuser 320may incorporate diffuser vanes. The diffuser vanes 325 may be used toimprove distribution of fluid flow within the diffuser 320. Similar tothe stationary pre-swirl vanes 323, the vanes within the diffuser and/orvolute may impart additional friction to the blood.

The embodiments disclosed herein may be designed with considerationsfrom the following references in mind, each of which is herebyincorporated by reference in its entirety. Considerations for geometricoptimization of centrifugal impellers related to MCSD specifications ofpressure rise, flow rate, diameter and rotational speed are describedby: Korakianitis, T., Rezaienia, M. A., Paul, G. M., Rahideh, A.,Rothman, M. T., Mozafari, S., “Optimization of Centrifugal PumpCharacteristic Dimensions for Mechanical Circulatory Support Devices”(2016) ASAIO Journal, 62 (5), pp. 545-551; and Mozafari, S., Rezaienia,M. A., Paul, G. M., Rothman, M. T., Wen, P., Korakianitis, T., “TheEffect of Geometry on the Efficiency and Hemolysis of CentrifugalImplantable Blood Pumps” (2017) ASAIO Journal, 63 (1), pp. 53-59.

The machinability of centrifugal impellers is described by: Paul, G.,Rezaienia, A., Avital, E., Korakianitis, T., “Machinability andoptimization of shrouded centrifugal impellers for implantable bloodpumps” (2017) Journal of Medical Devices, Transactions of the ASME, 11(2), art. no. 021005. The effects of a patient's motion on deviceoperation are described by: Paul, G., Rezaienia, A., Shen, X., Avital,E., Korakianitis, T., “Slip and turbulence phenomena in journal bearingswith application to implantable rotary blood pumps” (2016) TribologyInternational, 104, pp. 157-165; and Paul, G., Rezaienia, M. A.,Rahideh, A., Munjiza, A., Korakianitis, T., “The Effects of AmbulatoryAccelerations on the Stability of a Magnetically Suspended Impeller foran Implantable Blood Pump” (2016) Artificial Organs, 40 (9), pp.867-876.

The effects of device implantation in the descending aorta are describedby Rezaienia, M. A., Paul, G., Avital, E. J., Mozafari, S., Rothman, M.,Korakianitis, T. “In-vitro investigation of the hemodynamic responses ofthe cerebral, coronary and renal circulations with a rotary blood pumpinstalled in the descending aorta” (2017) Medical Engineering andPhysics, 40, pp. 2-10; Rezaienia, M. A., Paul, G., Avital, E., Rahideh,A., Rothman, M. T., Korakianitis, T., “In-vitro investigation ofcerebral-perfusion effects of a rotary blood pump installed in thedescending aorta” (2016) Journal of Biomechanics, 49 (9), pp. 1865-1872;Rezaienia, M. A., Rahideh, A., Alhosseini Hamedani, B., Bosak, D. E. M.,Zustiak, S., Korakianitis, T., “Numerical and In Vitro Investigation ofa Novel Mechanical Circulatory Support Device Installed in theDescending Aorta” (2015) Artificial Organs, 39 (6), pp. 502-513; andRezaienia, M. A., Rahideh, A., Rothman, M. T., Sell, S. A., Mitchell,K., Korakianitis, T., “In vitro comparison of two different mechanicalcirculatory support devices installed in series and in parallel” (2014)Artificial Organs, 38 (9), pp. 800-809.

Considerations for MCSD electric motor design are described by: Rahideh,A., Mardaneh, M., Korakianitis, T., “Analytical 2-D calculations oftorque, inductance, and back-EMF for brushless slotless machines withsurface inset magnets” (2013) IEEE Transactions on Magnetics, 49 (8),art. no. 6418033, pp. 4873-4884; Rahideh, A., Korakianitis, T.,“Analytical calculation of open-circuit magnetic field distribution ofslotless brushless PM machines” (2013) International Journal ofElectrical Power and Energy Systems, 44 (1), pp. 99-114; Rahideh, A.,Korakianitis, T., “Analytical magnetic field distribution of slotlessbrushless PM motors. Part 2: Open-circuit field and torque calculations”(2012) IET Electric Power Applications, 6 (9), pp. 639-651; Rahideh, A.,Korakianitis, T., “Analytical magnetic field distribution of slotlessbrushless permanent magnet motors—Part I. Armature reaction field,inductance and rotor eddy current loss calculations” (2012) IET ElectricPower Applications, 6 (9), pp. 628-638; Rahideh, A., Korakianitis, T.,“Analytical magnetic field calculation of slotted brushlesspermanent-magnet machines with surface inset magnets” (2012) IEEETransactions on Magnetics, 48 (10), art. no. 6203591, pp. 2633-2649;Rahideh, A., Korakianitis, T., “Subdomain Analytical Magnetic FieldPrediction of Slotted Brushless Machines with Surface Mounted Magnets”(2012) International Review of Electrical Engineering, 7 (2), pp.3891-3909; Rahideh, A., Korakianitis, T., “Analytical armature reactionfield distribution of slotless brushless machines with inset permanentmagnets” (2012) IEEE Transactions on Magnetics, 48 (7), art. no.6126045, pp. 2178-2191; Rahideh, A., Korakianitis, T., “Brushless DCmotor design using harmony search optimization” (2012) Proceedings—20112nd International Conference on Control, Instrumentation and Automation,ICCIA 2011, art. no. 6356628, pp. 44-50; Rahideh, A., Korakianitis, T.,“Analytical open-circuit magnetic field distribution of slotlessbrushless permanent-magnet machines with rotor eccentricity” (2011) IEEETransactions on Magnetics, 47 (12), art. no. 5893946, pp. 4791-4808;Rahideh, A., Korakianitis, T., “Analytical magnetic field distributionof slotless brushless machines with inset permanent magnets” (2011) IEEETransactions on Magnetics, 47 (6 PART 2), art. no. 5706366, pp.1763-1774; and Rahideh, A., Korakianitis, T., Ruiz, P., Keeble, T.,Rothman, M. T., “Optimal brushless DC motor design using geneticalgorithms” (2010) Journal of Magnetism and Magnetic Materials, 322(22), pp. 3680-3687.

Numerical simulations of the cardiovascular system with implanted MCSDsare described by: Shi, Y., Korakianitis, T., Bowles, C., “Numericalsimulation of cardiovascular dynamics with different types of VADassistance” (2007) Journal of Biomechanics, 40 (13), pp. 2919-2933;Korakianitis, T., Shi, Y., “Numerical comparison of hemodynamics withatrium to aorta and ventricular apex to aorta VAD support” (2007) ASAIOJournal, 53 (5), pp. 537-548; Shi, Y., Korakianitis, T., “Numericalsimulation of cardiovascular dynamics with left heart failure andin-series pulsatile ventricular assist device” (2006) Artificial Organs,30 (12), pp. 929-948; Korakianitis, T., Shi, Y., “Effects of atrialcontraction, atrioventricular interaction and heart valve dynamics onhuman cardiovascular system response” (2006) Medical Engineering andPhysics, 28 (8), pp. 762-779; Korakianitis, T., Shi, Y., “A concentratedparameter model for the human cardiovascular system including heartvalve dynamics and atrioventricular interaction” (2006) MedicalEngineering and Physics, 28 (7), pp. 613-628; and Korakianitis, T., Shi,Y., “Numerical simulation of cardiovascular dynamics with healthy anddiseased heart valves” (2006) Journal of Biomechanics, 39 (11), pp.1964-1982.

Devices for emulating the human cardiovascular system for in-vitrotesting of VADs and MCSD are described by: Ruiz, P., Rezaienia, M. A.,Rahideh, A., Keeble, T. R., Rothman, M. T., Korakianitis, T., “In vitrocardiovascular system emulator (Bioreactor) for the simulation of normaland diseased conditions with and without mechanical circulatory support”(2013) Artificial Organs, 37 (6), pp. 549-560.

Although the present invention has been described in terms of certainpreferred embodiments, it may be incorporated into other embodiments bypersons of skill in the art in view of the disclosure herein. The scopeof the invention is therefore not intended to be limited by the specificembodiments disclosed herein, but is intended to be defined by the fullscope of the following claims.

What is claimed is:
 1. A mechanical circulatory support for assistingthe heart, the support comprising: a casing comprising a main body, aninlet configured to introduce blood flow from an upstream portion of ahuman aorta into the main body, and an outlet configured to return theblood flow from the main body to a downstream portion of the humanaorta; an impeller positioned within an internal volume of the main bodyof the casing so as to receive blood flow from the inlet, the directionof the received blood flow defining a longitudinal axis, wherein theimpeller comprises a plurality of blades for pumping blood, the bladesbeing arranged around the longitudinal axis so as to define an outercircumference, and wherein the impeller is configured to rotate aroundthe longitudinal axis to pump the blood in a centrifugal manner towardthe outer circumference; and a diffuser integral with or joined to thecasing, the diffuser configured to receive blood outflow from theimpeller and direct the blood flow to the outlet, wherein the diffuseris at least partially open the internal volume of the main body of thecasing along at least a portion of the outer circumference of theimpeller.
 2. The support of claim 1, wherein the impeller is a shroudedimpeller comprising: an blade passage chamber; an upper portion forminga ceiling to the blade passage chamber, the upper portion having anupper channel extending along the longitudinal axis from a top of theimpeller to the blade passage chamber; and a lower portion forming afloor to the blade passage chamber, the lower portion having a lowerchannel extending along the longitudinal axis from the bottom of theimpeller to the blade passage chamber, wherein the blades extend from aninner circumference around the longitudinal axis to the outercircumference, the blades extending axially between the floor and theceiling of the blade passage chamber to join the upper portion and thelower portion together.
 3. The support of claim 2, wherein the casingfurther comprises a projection extending from the bottom of the casinginto the lower channel, and wherein the casing is configured to allowblood to flow from the outer circumference of the blades along secondaryflow paths between an internal surface of the casing and the lowerportion of the impeller, and between the projection and an internalsurface of the lower channel back to the blade passage chamber so as toprevent blood stagnation.
 4. The support of claim 1, wherein theimpeller is an unshrouded impeller.
 5. The support of claim 1, whereinthe impeller is magnetically suspended in an axial direction within thecasing by a combination of axial-suspension permanent magnets coupled toa top half and a bottom half of the casing and permanent magnetspositioned coupled to a top half and a bottom half of the impeller, theaxial-suspension permanent magnets coupled to the top half of the casingbeing axially spaced apart from the permanent magnets coupled to the tophalf of the impeller and the axial-suspension permanent magnets coupledto the bottom half of the casing being axially spaced apart from thepermanent magnets coupled to the bottom half of the impeller, andwherein the impeller is magnetically suspended in a radial directionwithin the casing by a radial-suspension permanent magnet coupled to thecasing near the permanent magnet in the top half of the impeller and bya radial-suspension permanent magnet coupled to the casing near thepermanent magnet in the bottom half of the impeller.
 6. The support ofclaim 5, wherein the impeller is configured to be radially stabilized byan eccentric hydrodynamic journal bearing force between the impeller andthe casing.
 7. The support of claim 5, wherein the impeller isconfigured to be radially stabilized by at least two electromagnetspositioned on opposite sides of each of the radial suspension permanentmagnets, wherein the force of each of the electromagnets is drivenaccording to impeller positioning information attained from eddy currentsensors coupled to the casing.
 8. The support of claim 7, wherein atleast one of the electromagnets coupled to the upper half of the casingis axially displaced from the permanent magnet coupled to the upper halfof the impeller and at least one of the electromagnets coupled to thelower half of the casing is axially displaced from the permanent magnetcoupled to the lower half of the impeller, and wherein the position ofthe impeller is configured to be oscillated in the axial direction tocreate a pulsatile flow by pulsatile phases of current applied to theelectromagnets.
 9. The support of claim 1, further comprising a motorfor electromagnetically rotating the impeller around the axialdirection, the motor comprising: a stator within the casing comprising aplurality of electromagnets; and a rotor within the impeller comprisinga plurality of permanent drive magnets, the rotor configured to bepositioned concentrically within the stator.
 10. The support of claim 1,wherein the support is configured to create a vortex in an outflow ofblood exiting the outlet to emulate the naturally-occurring vortex inthe native aorta of a healthy human heart.
 11. The support of claim 1,wherein the support is configured to create a pressure rise in theintroduced blood flow between about 40 mmHg and about 80 mmHg andwherein the support is configured to maintain a blood flow rate of about5 L/min.
 12. The support of claim 1, wherein the support is configuredto be installed in-series with a portion of the descending aorta of ahuman aorta.
 13. The support of claim 1, wherein the inlet is configuredto redirect the blood flow 90 degrees before it enters the main body,such that the inlet and the outlet are parallel with each other.
 14. Thesupport of claim 1, wherein the blood flow is redirected toward an axialdirection prior to reaching the outlet, such that the outlet issubstantially collinear with the inlet.
 15. The support of claim 14,wherein the diffuser wraps around the casing in a spiral configurationto facilitate the formation of a vortex in the outflow which emulatesthe naturally-occurring vortex in the native aorta of a healthy humanheart.
 16. The support of claim 14, further comprising a splitter vanepositioned within at least a portion of the diffuser which rotates withrespect to a circumference of the diffuser to facilitate the formationof a vortex in the outflow which emulates the naturally-occurring vortexin the native aorta of a healthy human heart.
 17. The support of claim14, further comprising a splitter vane positioned within at least aportion of a volute of the outlet which rotates with respect to acircumference of the volute to facilitate the formation of a vortex inthe outflow which emulates the naturally-occurring vortex in the nativeaorta of a healthy human heart.
 18. The support of claim 14, furthercomprising a plurality of diffuser vanes positioned circumferentiallyaround the outer circumference defined by the impeller.
 19. The supportof claim 14, further comprising a plurality of stationary pre-swirlvanes positioned within in inlet.
 20. The support of claim 1, wherein aportion of a surface of the internal volume of the main body of thecasing and/or a portion of an outer surface of the impeller comprisesspiraling grooves configured to facilitate secondary flow paths of bloodbetween the impeller and the casing.
 21. A method of treating congestiveheart failure in a patient, the method comprising: installing amechanical circulation support within the descending aorta of thepatient, wherein the mechanical circulation support comprises acentrifugal blood pump configured to provide a pressure rise betweenabout 40 mmHg and about 80 mmHg in the blood flow and to maintain a flowrate of about 5 L/min.
 22. The method of claim 21, wherein the supportis installed in series with the descending aorta, the method furthercomprising severing the aorta into upper and lower portions, wherein theinstalling comprises grafting the upper portion to an inlet of thesupport and grafting the lower portion to an outlet of the support 23.The method of claim 21, wherein the support is installed in parallelwith the descending aorta, the method further comprising installing aone-way valve in the native aorta in parallel with the support, suchthat blood cannot flow upstream through the native aorta to recirculatethrough the support.
 24. The method of claim 21, wherein the support isinstalled such that both an inlet to the support and an outlet from thesupport are oriented at a non-linear angle to the native aorta.
 25. Themethod of claim 21, wherein the support is installed such that both aninlet to the support and an outlet from the support are oriented to besubstantially collinear with the native aorta.
 26. The method of claim21, wherein the support is installed such that both an inlet to thesupport and an outlet from the support are oriented to be parallel withthe native aorta.
 27. The method of claim 21, wherein the patient hasstage III or stage IV congestive heart failure.
 28. The method of claim21, wherein the patient has late stage III or early stage IV congestiveheart failure.